Authors: Kun Wu, Zhihe Yun, Wu Xue, Tao Yu, Anyuan Dai, Inbo Han, Vit Kotheeranurak, Worawat Limthongkul, Yanting Liu, Qinyi Liu
Categories: Review Article, injectable hydrogels, Neural tissue engineering, Central nervous system, Peripheral nervous system, Regenerative medicine
Source: Materials Today Bio
Authors: Kun Wu, Zhihe Yun, Wu Xue, Tao Yu, Anyuan Dai, Inbo Han, Vit Kotheeranurak, Worawat Limthongkul, Yanting Liu, Qinyi Liu
The nervous system, comprising the central and peripheral components, is essential for regulating physiological functions. Damage to neural tissue often results in severe motor and sensory deficits, with limited endogenous regenerative capacity posing major challenges for repair. Tissue engineering offers promising strategies for neural regeneration by integrating cells, biomaterial scaffolds, and drug delivery systems. Among biomaterials, injectable hydrogels have attracted significant attention due to their biocompatibility, biodegradability, flexibility, and ability to undergo minimally invasive in situ gelation. These hydrogels serve as versatile carriers for cells and therapeutic agents, enabling precise delivery to hard-to-reach tissues while promoting neural repair. This review highlights recent advances in the design, preparation, and application of injectable hydrogels for neural tissue engineering. We discuss their roles in regenerating the brain, spinal cord, and peripheral nerves, address current challenges, and provide perspectives on future developments. Collectively, injectable hydrogels represent a promising platform for developing next-generation regenerative therapies aimed at restoring both anatomical structure and functional outcomes in neural tissues.
The nervous system is a cornerstone of human physiology, representing the most intricate network among all bodily systems [1]. It encompasses the central nervous system (CNS) and the peripheral nervous system (PNS). The central nervous system consists of the brain and spinal cord, which contain a dense network of neurons and supporting glial cells. Functionally, the CNS and PNS collaborate to form neural circuits that regulate vital activities and facilitate the transmission of physiological signals [2,3]. Damage or injury to the nervous system can lead to significant motor and sensory deficits, profoundly impacting the patient quality of life [4].
Both the central and peripheral nervous systems face formidable repair challenges due to the intricate architecture of neural tissue and its limited inherent regenerative capacity. Current clinical treatment interventions primarily include surgery and pharmacotherapy; however, often fall short in promoting effective neural repair [5]. Tissue engineering, a multidisciplinary field, involves developing biological substitutes to restore, regenerate, or replace defective tissues using a combination of cells, biomaterial scaffolds, and drug delivery systems [6]. To better address research and clinical needs, tissue engineering has expanded into various biomedical fields, it has enhanced diagnostic tools and opened new avenues for therapeutic interventions [7]. Consequently, achieving both anatomical and functional regeneration of neural tissue has become a key topic of interest in tissue engineering. Restoring damaged neural tissue both functionally and anatomically, with advances in biomaterials paving the way for innovative regenerative therapies aimed at restoring neural function [8,9].
Hydrogels are three-dimensional soft materials characterized by a highly porous structure, capable of retaining large amounts of water and exhibiting high hydrophilicity. They also exhibit high biocompatibility, biodegradability, solid elasticity, deformability, and flexibility [10]. These unique properties establish hydrogels as a promising platform for biological materials [11], with applications in various biomedical fields, including tissue engineering, wound dressings, drug delivery, and cell carriers [12,13]. Among the different types available, injectable hydrogels are particularly attractive due to their ease of administration and ability to form gels in situ [14,15](Fig. 1). Furthermore, as implants, injectable hydrogels provide less invasive delivery than traditional approaches, enabling access to the tissues which are hard-to-reach while simultaneously serving as carriers for drugs or cells [13].Fig. 1Preparation of injectable hydrogels and application in central and peripheral nervous system.Fig. 1
Grimaudo et al. [16] reviewed the applications of injectable hydrogels in the CNS, highlighting recent innovations and insights into neural tissue responses. However, their work did not offer a comprehensive analysis of the effects of different injectable hydrogels on both the spinal cord and brain. Almawash et al. [12] conducted an in-depth review of recent developments in several key types of injectable hydrogels, emphasizing their preparation. Yet, their depiction of injectable hydrogels across various systems was overly broad and lacked detailed descriptions of their applications in the nervous system. George et al. [14] examined potential materials for CNS tissue engineering, addressing material selection and hydrogel structure customization, but offered only minimal discussion of injectable hydrogels. Hasanzadeh et al. [17] explored the pathophysiology of CNS disorders alongside the applications of various injectable hydrogels. Given the differing requirements for biomaterials in specific brain disorders, describing brain tissue as a whole may not be the most effective approach. Notably, none of the aforementioned reviews addressed the applications of injectable hydrogels in the PNS or clarified the relationship between their roles in the central and peripheral nervous systems [18].
In contrast, this review focuses on injectable hydrogels in neural tissue engineering, we emphasizing various design principles and preparation strategies. We further emphasize the current applications of injectable hydrogels in the regeneration of brain, spinal cord, and peripheral nerve tissues. Finally, we discuss the development, potential applications, challenges, and future directions of injectable hydrogels in neural tissue engineering.
Following the overview of neural tissue engineering approaches and the applications of injectable hydrogels in the CNS and PNS, it is essential to understand the anatomical and cellular foundations of neural tissue. Functionally, the nervous system is primarily composed of nervous tissue, which contains two principal cell neurons and glial cells. Neurons are the fundamental structural and functional units responsible for transmitting electrical and chemical signals. Each neuron consists of a cell body (soma), dendrites, and an axon. The cell body contains a large nucleus with a prominent nucleolus, serving as the metabolic center of the neuron. Dendrites receive incoming signals, whereas the axon transmits electrical impulses to target cells. Glial cells, including astrocytes and oligodendrocytes in the 10.13039/100015862CNS and Schwann cells in the PNS, provide essential metabolic and structural support to neurons [8].
The PNS is considered the simpler division of the nervous system and includes all neural structures outside the CNS. It consists primarily of afferent sensory fibers, which convey information to the CNS, and efferent motor fibers, which transmit commands to peripheral effectors through neuromuscular junctions [19]. Bundles of these nerve fibers are encased in connective tissue sheaths, forming distinct peripheral nerves that appear as whitish, cord-like structures. Anatomically, the PNS is subdivided into cranial nerves, spinal nerves, and autonomic nerves. Cranial nerves connect directly to the brain, spinal nerves emerge from the spinal cord, and autonomic nerves regulate involuntary functions through the sympathetic and parasympathetic systems.
In the CNS, regions rich in neuronal cell bodies and dendrites are known as gray matter, which forms the outer layer of the brain (the cortex) and the inner core of the spinal cord. Conversely, regions with dense axonal tracts are called white matter, appearing as the medulla in deeper brain layers. In cross-sectional views of the spinal cord, gray matter occupies the central region, shaped like a butterfly, surrounded by white matter. The dorsal horn contains interneurons that receive sensory input, whereas the ventral horn houses somatic motor neurons that project to skeletal muscles, mediating both voluntary and reflexive movements.
The brain serves as the command center, integrating sensory input and coordinating motor output, emotions, communication, and cognition. The spinal cord acts as a bidirectional communication highway—transmitting motor signals from the brain to the periphery and sensory information back to the brain. The PNS complements this system by conveying sensory data to the CNS via afferent fibers and delivering motor commands from the CNS via efferent fibers [20]. Together, the CNS and PNS maintain the body's homeostasis and enable complex physiological and behavioral functions.
Building upon the anatomical and structural features of the CNS and PNS described, understanding the pathophysiological responses following neural injury is critical for developing effective repair strategies. Regeneration following peripheral nerve injury relies on the coordinated activity of various cell types (Fig. 2a). Following injuries such as nerve compression, contusion, or complete transection, the distal segment of the peripheral nerve initiates a process called Wallerian degeneration [21]. Initially, the severed axonal segments undergo cytoskeletal disintegration and eventually breaking down into cellular debris. Over time, Schwann cells, essential for peripheral nerve repair, become activated. These cells undergo dedifferentiation, degrade their myelin sheaths, and phagocytose extracellular debris [22,23]. During this process, some Schwann cells temporarily lose their normal myelination capacity and appear as abnormal myelinating Schwann cells, while others remain non-myelinating Schwann cells that ensheath small axons without forming compact myelin. Both subtypes contribute to debris clearance and axonal guidance in the regenerative phase. Activated schwann cells further secrete growth factors that promote axonal growth and support neuronal survival while also releasing cytokines and chemokines to recruit immune cells to the damaged nerve. Recruited immune cells, primarily macrophages, play a crucial role in clearing debris and producing growth factors during Wallerian degeneration [24,25]. In essence, peripheral nerve tissue continuously regenerates within a complex, multicellular cooperative environment and effective repair hinges on enhancing the collaborative response of these cell types to create a favorable regenerative environment for damaged nerve cells, ultimately facilitating the recovery of nerve tissue [26].Fig. 2Schematic illustrations of the pathophysiology and barrier mechanisms in the nervous system. (a) Cellular events during PNS injury, emphasizing Wallerian degeneration and Schwann cell–mediated regeneration. (b) Key cellular and molecular responses following CNS injury, highlighting limited regenerative capacity and glial scar formation. (c) Architecture and protective role of the BNB, showing how peripheral nerves are shielded while permitting selective molecular exchange. (d) Structural components and selective permeability of the BBB, illustrating challenges for therapeutic delivery in the CNS.Fig. 2
In contrast to the robust, coordinated response seen in peripheral nerve injuries, the intrinsic healing capacity of endogenous cells in the CNS is severely limited by both cellular and extracellular environmental factors (Fig. 2b), resulting in negligible axonal regeneration and functional recovery [27].
First, CNS inflammation, driven by persistent pro-inflammatory macrophage responses, exhibits greater cytotoxicity than that observed in peripheral nerve. Resting microglia serve as immune sentinels, maintaining homeostasis and clearing minor debris. After injury, these cells transform into activated microglia, releasing pro-inflammatory cytokines and amplifying the inflammatory cascade. Moreover, macrophages in the CNS are less efficient at clearing cellular debris, leading to the accumulation at the injury site, and further impeding nerve healing and repair [28,29]. Second, the composition of the extracellular matrix (ECM) in the CNS differs significantly from that in the PNS. The CNS ECM has a high concentration of glycosaminoglycans, such as hyaluronic acid, glycoproteins such as tenascin-C, and platelet-derived growth factors, yet it lacks adequate amounts of collagen, laminin, and fibronectin, which are essential for providing tissue structure and strength [24]. Finally, although Schwann cells are central to peripheral nerve repair, astrocytes serve as the primary repair cells in CNS. Oligodendrocytes, responsible for myelinating CNS axons, often undergo apoptosis or enter a quiescent state following axonal injury. In comparison to Schwann cells, oligodendrocytes are more vulnerable to damage, offer limited support for axonal growth, and exhibit minimal phagocytic activity. Astrocytes, in response to CNS injury through proliferation, hypertrophy, and upregulation of glial fibrillary acidic protein (GFAP), which contributes to the formation of "glial scars". Although these scars are generally regarded as detrimental to nerve function recovery, these scars also serve as physical and molecular barriers to axonal regeneration by sealing the injury site, preventing further damage, modulating blood flow, and influencing immune responses. [24,30]. Thus, scar tissue following CNS injury presents a complex role that cannot be easily classified as solely beneficial or harmful to CNS repair.
Generally, PNS possess superior regenerative capabilities compared to the CNS. However, even the PNS has superior regenerative abilities, the responses of injured neurons and surrounding cells are often not sustained long enough to promote robust growth, leading to less-than-optimal repair outcomes. In this CNS, the complex post-injury environment further complicates recovery, where key challenges include managing persistent inflammation, optimizing the composition of the ECM, and harnessing the potentially protective aspects of scar tissue.
Injectable hydrogels offer a promising solution for addressing these challenges in both systems. In PNS injuries, these hydrogels can maintain growth-promoting responses over extended periods. Thereby enhancing tissue regeneration. In 10.13039/100015862CNS injuries, injectable hydrogels help establish an environment conducive to recovery by mitigating adverse inflammatory responses and providing structural support.
Understanding the anatomical structures and pathophysiological responses of the CNS and PNS provides essential context for appreciating the unique barrier mechanisms that regulate their internal environments. The CNS is protected from external insults by several specialized barriers, including the blood-brain barrier (BBB), the blood-cerebrospinal fluid barrier (BCSFB), and the blood-spinal cord barrier (BSCB) [[31], [32], [33]]. The BBB serves as a natural barrier, separating the brain from systemic circulation [[34], [35], [36]] (Fig. 2d). Its primary functions include preventing the entry of toxic substances, protecting the brain, and maintaining a stable microenvironment within the CNS [37,38]. Endothelial cells (ECs) form the primary cellular components of the BBB [39], while tight junction complexes, pericytes, astrocytes, and microglia collectively reinforce its structure integrity [34,38]. However, this selective barrier also restricts the entry of many therapeutic molecules [40]. For instance, brain microvascular endothelial cells (BMECs) express ATP-binding cassette (ABC) family efflux transporters that primarily remove metabolic waste and potential toxins but these transporters also expel therapeutic drugs, thereby reducing their efficacy [41,42]. Additionally, tight junction complexes and associated glial cells limit passive paracellular diffusion [43]. Consequently, large therapeutic drugs and approximately 98 % of small molecules cannot penetrate the BBB. Only small lipophilic molecules (<500 Da) can cross in effective concentrations, significantly limiting the application of therapeutic drugs for treating brain diseases [44].
The BSCB functions similarly to the BBB in regulating molecular movement between the blood and spinal cord parenchyma, exhibits distinct features Recently, disruptions in BSCB integrity are increasingly recognized as pivotal contributors to the pathogenesis and progression of spinal cord disorders, including traumatic spinal cord injury. Such dysfunctions can initiate and exacerbate inflammatory cascades and secondary injury mechanisms, thereby further compromising spinal cord tissue and impeding intrinsic regenerative processes [45]. Its integrity depends on non-fenestrated endothelial cells in spinal cord capillaries, as well as supporting components such as the basement membrane, pericytes, and astrocytic end-feet [46]. The coordinated interactions among these elements are essential for maintaining the barrier's protective functions. In contrast, the BCSFB separates cerebrospinal fluid (CSF) within the brain's ventricles from the bloodstream and plays a pivotal role in mediating the inflammatory response between the peripheral and central nervous systems [47,48]. The BCSFB is located in the choroid plexus (CP) of the ventricles, where the integrity of the CP is crucial for maintaining brain homeostasis and BCSFB permeability, as it is a major site for CSF secretion [49]. The BCSFB consists of various cell junctions CP epithelial cells. Tight junctions at the apical surfaces of these cells regulate the permeability and integrity by controlling the paracellular diffusion of water-soluble molecules [32]. Moreover, active organic acid transporters in the choroid plexus enhance the BCSFB by exporting organic acids from the CSF into the blood. resulting in the rapid clearance of many organic acid-based therapeutics [50,51].
In the PNS, several protective structures work synergistically to shield neural elements from external damage. These include the endoneurium, a collagenous network surrounding each axon; the perineurium, which encircles axon bundles; and the epineurium, composed of longitudinally arranged collagen fibers [52]. Recent research indicates that the perineurium and endoneurial vasculature are essential components of the blood-nerve barrier (BNB) [43] (Fig. 2c). The BNB regulates the exchange between the nerve microenvironment and systemic circulation, restricting the infiltration of soluble mediators and leukocytes into peripheral nerve tissue [53]. Analogous to BBB dysfunction in CNS disorders, impairment of the BNB can disrupt the endoneurium homeostasis, facilitating the entry of toxic substances that may precipitate peripheral neuropathies and further neural deterioration [53]. Thus, BNB protection of the endoneurium is essential for maintaining normal PNS function. Moreover, Schwann cells in the PNS synthesize a lipid substance known as myelin, which wraps concentrically around axons to form the myelin sheath. This process not only insulates and segregates individual neurons but also significantly enhances conduction velocity, thereby ensuring rapid and efficient neural signal transmission.
The barriers in the CNS protect the brain and spinal cord while posing challenges for delivering therapeutic agents across them [54]. Traditional methods for delivering therapeutics directly to CNS target sites typically rely on surgical approaches, including intrathecal drug delivery and convection-enhanced delivery [44]. However, these methods are often associated with risks to patient safety and and may result in suboptimal drug concentrations at target sites, thereby limiting therapeutic efficacy. Similarly, addressing peripheral nerve injuries in the PNS often requires the direct delivery of targeted drugs or therapeutic factors to the affected nerve [43,55]. Although intra-nerve injections enable localized delivery, they carry risks such as mechanical disruption of the nerve barrier and potential axonal damage due to uncontrolled high-flow infusions of extracellular matrix components combined with therapeutic agents [56,57]. Injectable hydrogels offer a promising solution to these challenges in both the central and peripheral nervous systems. By providing a minimally invasive delivery route, they can establish a sustained, localized environment that enables gradual release of therapeutic molecules, thereby enhancing treatment efficacy while mitigating adverse effects associated with conventional delivery methods. Recent advances in hydrogel engineering have further enabled these materials to help overcome neural barriers through multiple mechanisms. Physically, injectable hydrogels can penetrate regions where the BBB or BSCB is compromised after injury, forming an in situ depot for controlled and sustained release. Chemically, functionalization with cell-penetrating peptides, targeting ligands (e.g., transferrin, lactoferrin), or nanoparticle conjugation facilitates transcytosis across endothelial cells. Moreover, stimuli-responsive hydrogels that react to pH, temperature, or enzymatic cues allow precise drug release within the microenvironment [58]. In peripheral nerve repair, hydrogel matrices can transiently modulate blood–nerve barrier permeability and provide extracellular matrix–like support that promotes Schwann cell migration and axonal regeneration [59].
Considering the complex architecture of the nervous system, the pathophysiological mechanisms following injury, and the existing barrier challenges, the rational design of injectable hydrogels to meet mechanical, structural, and biological requirements is crucial for achieving effective neural repair. Injectable hydrogels, like other biomaterials used in tissue engineering, share common characteristics such as biocompatibility, non-immunogenicity, non-toxicity, and biodegradability. Beyond these fundamental characteristics, injectable hydrogels possess unique attributes, specialized rheological behaviors, including shear-thinning and self-healing, which facilitate in situ gelation and effective delivery through standard needles (e.g., 18G–26G) [60]. Moreover, their production must adhere to stringent control standards to minimize harmful or excessive byproduct generation.
For neural tissue engineering applications, injectable hydrogels must meet several critical design criteria [61]. Firstly, injectable hydrogels must provide sufficient mechanical support upon reaching the target site. Neural tissue scaffolds encounter mechanical stresses from the surrounding tissues [62]. Thus, the hydrogel's storage modulus should exceed its loss modulus to ensure a stable, gel-like state that resists collapse under tissue load [63]. Secondly, the pore architecture of the hydrogel is paramount. It is crucial for cell adhesion and migration, regulates the diffusion of media and metabolic products, and ultimately supports cell survival [64]. Research indicates that an optimal pore diameter of 10–100 μm, coupled with a scaffold porosity of approximately 90 %, is ideal for fostering neural cell growth, promoting tissue differentiation, and sustaining cell viability [65,66]. Degradation kinetics of injectable hydrogels is another critical characteristic that needs to be controlled applied to neural tissue engineering An overly rapid degradation rate can lead to glial scar formation, induce inflammation, and inhibit neural regeneration, each of which can impede neural regeneration and cause swelling at the injection site [67]. Thus, a balanced degradation profile is essential to maintain the scaffold's structural integrity during the critical phases of tissue repair. Ultimately, the primary goal of employing injectable hydrogels in neural tissue engineering is to facilitate effective nerve repair and functional restoration [68]. Given the heightened vulnerability of neural cells compared to other cell types, the design of these hydrogels necessitates rigorous specifications to ensure both safety and therapeutic efficacy. Notably, the viscoelastic properties of hydrogels play a critical role in regulating neural stem cell migration, proliferation, and differentiation. Evidence indicates that hydrogels with appropriately tuned viscoelasticity can enhance NSC growth and promote axonal extension. Specifically, hydrogels exhibiting moderate viscoelasticity provide an extracellular matrix–mimetic microenvironment that supports neural network formation [69]. Finally, the mechanical properties of neural tissues vary across different regions. Specifically, the elastic modulus of brain tissue generally ranges from 0.1 to 1.5 kPa [70,71], while the spinal cord exhibits slightly higher stiffness values of approximately 0.5–3 kPa [72,73]. In contrast, peripheral nerve tissues typically present moduli between 5 and 50 kPa [74,75], depending on the anatomical location and testing conditions. Therefore, tailoring hydrogel stiffness within these ranges is essential to achieve biomechanical compatibility and support appropriate neural cell behavior.
Building on these design standards, designing and preparing injectable hydrogels with exceptional performance for neural tissue regeneration are crucial for their future translation into clinical applications. Thus, selecting hydrogel materials and the methods for their preparation and synthesis are critical to developing injectable hydrogels with translational potential.
An ideal tissue engineering scaffold must possess low toxicity, excellent biocompatibility, and biodegradability, while also providing solid support for cell proliferation. Additionally, it should promote cell differentiation and new tissue formation, all within a degradation profile that aligns with the tissue regeneration timeline [76,77]. Injectable hydrogels intended for neural tissue engineering are especially required to meet these rigorous criteria. To date, a diverse array of biomaterials—both natural and synthetic—has been employed to fabricate injectable hydrogel scaffolds, each offering unique advantages and challenges in the context of neural repair.
Natural polymers are typically classified into three protein-based, polysaccharide-based, and nucleotide-based. Compared to synthetic polymers, natural polymers demonstrate superior biocompatibility, greater ease of degradation and remodeling by cells, and involvement in biological signaling pathways [78,79]. However, they also have drawbacks, including lower mechanical properties, potential immunogenicity, and significant variability in their properties due to differences in natural sources and batch-to-batch inconsistencies. Moreover, precise control over their chemical structure, molecular weight, and degradation behavior remains a substantial challenge [80,81].
Hyaluronic Acid (HA) is a high-molecular-weight, linear anionic polysaccharide that is a major component of the ECM of all mammals. It is particularly abundant in the brain, especially within the niche of neural progenitor cells (NPCs) [82]. HA demonstrates excellent biodegradability, biocompatibility, and low toxicity while its derivatives can trigger receptor-mediated signal transduction in cells [83,84]. In the presence of water, the reactive carboxyl and hydroxyl groups of HA facilitate its modification and crosslinking to form hydrogels, making it a widely applicable material for injectable hydrogel systems. Studies have modified HA to incorporate acrylate groups and utilized degradable crosslinkers, achieving gelation via Michael addition chemistry mediated by matrix metalloproteinases (MMPs). By adjusting the ratios of these crosslinkers, researchers have fabricated HA-based hydrogels with mechanical properties closely resembling those of brain tissue [85]. Recent research has focused on developing self-healing hydrogels through the semi-interpenetrating polymer network (SIPN) by blending HA with chitosan-based hydrogels. Finding indicate that an increased HA content leads to a denser nanoscale network and a more porous microstructure which in turn enhances in vitro diffusion, cellular proliferation, migration, and differentiation of neural stem cells. These HA-incorporated SIPN hydrogels not only demonstrate superior biocompatibility but also actively promote neural tissue repair, underscoring their potential in treating neural system injuries [86].
Chitosan is a linear polysaccharide derived from the deacetylation of chitin, commonly obtained from the exoskeletons of crustaceans. It is composed of β-(1,4)-linked D-glucosamine and N-acetyl-D-glucosamine units. Similar to hyaluronic acid, chitosan exhibits excellent biodegradability, biocompatibility, and low toxicity, along with intrinsic anti-inflammatory and antioxidant properties [[87], [88], [89]]. However, its poor solubility under physiological pH and limited mechanical strength restrict its practical applications.
To overcome these limitations, chemical modifications of chitosan have been extensively explored. Functionalization with carboxyl, hydroxyl, or phenolic groups can significantly enhance solubility, mechanical stability, and biological activity. For example, grafting ferulic acid or succinic acid onto the chitosan backbone improves both mechanical integrity and anti-inflammatory performance, facilitating neural tissue repair following CNS injury [90].
Among chitosan derivatives, carboxymethyl chitosan (CMC) has been the most extensively investigated. Compared with native chitosan, CMC exhibits superior water solubility, pH responsiveness, and mechanical stability. Numerous studies have demonstrated its potential for both CNS and PNS regeneration. For instance, Shubin Li et al. recently synthesized a catechol- and choline-functionalized CMC derivative, which showed promising neural repair potential and favorable biocompatibility in peripheral and central nerve models [91]. In addition, Shui Guan et al. reported a conductive CMC/gelatin/PEDOT composite hydrogel that significantly promoted neuronal differentiation and neural repair, serving as a conductive scaffold for nerve tissue engineering applications [92]. These studies underscore the versatility of CMC-based hydrogels as injectable scaffolds for neural regeneration.
The mechanical properties of hydrogels play a critical role in regulating neural stem cell (NSC) behavior. NSCs preferentially differentiate into neurons within softer matrices (0.1–1 kPa) and into glial cells in moderately stiff matrices (7–10 kPa). Accordingly, self-healing chitosan-based hydrogels with an optimized stiffness of approximately 1.5 kPa have been developed for CNS repair. Experimental evidence indicates that neural spheroid progenitor cells proliferate nearly twice as fast in these injectable hydrogels compared to conventional formulations, while also exhibiting enhanced neuronal differentiation potential, highlighting their promise for neural tissue engineering [93].
Collagen is the most abundant protein in the human body and plays a critical structural role by contributing to tissue mechanical properties and overall shape. It interacts with cells via multiple receptors, regulating their cellular proliferation, migration, and differentiation. Specific collagen types are distributed in a tissue-dependent manner, thereby playing essential roles in nervous system development and making collagen an excellent candidate for nerve tissue regeneration [94]. Recent studies highlight the therapeutic potential of collagen-based hydrogels in neural repair. For example, bone marrow-derived mesenchymal stem cells (MSCs) loaded into collagen hydrogels and injected into traumatic brain injury (TBI) models have demonstrated superior outcomes compared to MSCs alone. These collagen hydrogel-MSC composites not only enhanced in vivo cell survival and neurite growth but also improved neurological recovery, as evidenced by positron emission tomography (PET) assessments that revealed enhanced brain metabolism. Furthermore, researchers also employed injectable collagen hydrogel as a carrier to deliver GDNF-overexpressing MSCs. In vitro, finding showed that collagen hydrogel maintain cell viability and did not impede GDNF secretion into the surrounding medium, while in vivo, it facilitated GDNF secretion into striatal tissue, positively influencing nerve repair. Thereby, the injectable collagen-based hydrogel is a well-tolerated platform for cell delivery, offering significant potential for neural tissue repair and the treatment of brain injuries [95].
Fibrin is a natural fibrous protein essential for blood coagulation. It enhances cell adhesion, motility, proliferation, and differentiation [96]. In the context of injectable hydrogels, fibrin serves as an excellent matrix due to its biocompatibility, biodegradability, and strong tissue bonding capabilities. When derived from autologous sources, fibrin minimizes the risk of foreign body reactions, making it particularly attractive for clinical applications. It exhibits high flexibility, making it especially suitable for delicate nervous system tissues [97,98]. Some researchers have utilized fibrin hydrogels loaded with xenogeneic MSCs to treat spinal cord injuries (SCI). Researchers combined xenogeneic sheep bone marrow-derived MSCs with fibrin hydrogels and implanted this composite into the injured spinal cords of rats. The results indicated that the sheep MSCs were capable of traversing the BBB, reconstructing the glial cell population, and differentiating into neuron-like and microglia-like cells, thereby mitigating neural tissue damage following SCI [99]. Moreover, to further enhances the affinity of fibrin hydrogels for pluripotent stem cell-derived neural stem/progenitor cells (NSPCs) and to support their migration and neurite extension, researchers explored incorporating synthetic peptides that bind to integrin α6β1, a receptor prevalent in NSPCs. Ultimately, fibrin hydrogels functionalized with the synthetic peptide HYD1, which targets integrin α6β1, promoted neurite extension of Embryonic stem (ES)-derived neural stem/progenitor cells (ES-NSPCs) and improved functional recovery following complete spinal cord transection, suggesting new directions for CNS regeneration [100].
Silk fibroin has attracted increasing attention in neural tissue engineering due to its excellent mechanical strength, controlled degradability, and sustained release characteristics. Silk fibroin–based hydrogels not only provide robust physical support but also facilitate neural cell adhesion and axonal extension through their β-sheet structure [101,102]. In addition, ECM-derived proteins obtained from decellularized tissues—such as laminin and fibronectin—can mimic the native microenvironment, regulate cell differentiation, and modulate immune responses. These materials have demonstrated promising outcomes in spinal cord injury and traumatic brain injury repair in recent years [103,104]. Therefore, composite hydrogels combining silk fibroin and ECM components are considered a promising next-generation platform for neural tissue regeneration.
Synthetic polymers offer an alternative strategy for nerve tissue engineering. Currently, the primary synthetic injectable hydrogels utilized in neural tissue applications are polyethylene glycol (PEG)-based and polyacrylamide hydrogels. While synthetic materials offer advantages such as high reproducibility and tunable properties, they often require additional functionalization to improve cell-matrix interactions, their inherent physical properties are generally less favorable than those of natural polymers [105]. Moreover, they are more prone to induce inflammatory responses following in vivo administration [14].
Polyethylene Glycol (PEG) is a widely studied synthetic, biodegradable, water-soluble polyether in tissue engineering application due to its non-toxic, non-immunogenic nature, and resistance to contamination, particularly for articular cartilage and nerve tissue regeneration [106]. Research has shown that PEG hydrogels implanted in cavities facilitate the sustained migration of astrocytes, filling voids, reducing cell aggregation, and inhibiting scar formation [107,108]. This functional profile may account for PEG hydrogels' ability to rapidly restore nerve conduction after severe SCI, promoting axonal myelination and enhancing both sensory and motor functions [109]. Furthermore, researches into the neuroinflammatory responses examined that the local recruitment and activation of microglia and astrocytes after implanting PEG-based hydrogels in the brain. They also found that PEG-based hydrogels not only demonstrate compatibility with brain tissue but also reduce acute and chronic inflammatory responses induced by the implantation process.
Polyacrylamide (PAM) hydrogels have garnered considerable attention and in biomedicine research due to their tunable mechanical properties [110], making PAM-based hydrogels a preferred choice for studying cell-matrix mechanical interactions [111,112]. A study from Nic D. Leipzig's team found that methylacrylamide chitosan (MAC) hydrogels with a stiffness below 1 kPa can induce the differentiation of adult NSC [113]. In addition, due to its high biocompatibility and ability to modulate viscoelasticity, PAM is widely used in hydrogels for neural repair. [114]. Ales Hejcl et al. developed macroporous hydrogels based on 2-hydroxyethyl methacrylate (HEMA), which adhered effectively to spinal tissue in a rat spinal cord transection model. They found that this hydrogel can bridge the spinal cavity after injury and provide a scaffold for the ingrowth of regenerating axons [115]. However, despite these promising features, due to concerns regarding its toxicity and suboptimal cell adhesion properties, PAM is not as widely applied as PEG [110].
Gelation is a crucial step in preparing injectable hydrogels. In nerve tissue engineering, selecting an appropriate gelation method is essential, as it must align with the designed structure and intended application of the hydrogel. Broadly, gelation methods for injectable hydrogels are classified into two chemical and physical crosslinking reactions (Fig. 3).Fig. 3Chemical Crosslinking. Photo/light-responsive, Yantao Zheng et al. and Wanlin Xu et al. both used blue (405 nm) light to prepare gelatin methacryloyl (GelMA) hydrogel and applied them respectively to the central and peripheral nervous systems [150,151]. Click Chemistry, the Diels-Alder click-reaction between HA-furan and polyethylene glycol(PEG)-bismaleimide can result in hyaluronic acid (HA) hydrogel [152]. Enzymatically, Horse radish peroxidase (HRP) and choline oxidase (ChOx) has been used for gelatin-hydroxyphenyl (GH) hydrogel crosslinking [153], while hyaluronic acid hydrogel in situ dual-enzymatically was cross-linked by galactose oxidase (GalOx) and horseradish peroxidase [154]. Michael addition, Curcumin-loaded lysine/poly(ethylene glycol) diacrylate (PEGDA) hydrogels were synthesized via Michael addition for the treatment of traumatic brain injury [155]. Similarly, ibuprofen-KYIGSRK conjugated hydrogels were synthesized in situ via Michael addition, without the need for catalysts or additional reaction steps [156]. Schiff base, Hyaluronic acid-based hydrogels (AHA/DTP hydrogels) were obtained through dynamic Schiff base linkages between the dialdehyde groups of aldehyde-modified hyaluronic acid (AHA) and the hydrazide groups of 3, 3′-dithiobis (propionylhydrazide) (DTP) [157].Fig. 3
Chemical crosslinking involves the formation of covalent bonds to establish a stable, interconnected network. This process can be achieved through several approaches, including Michael addition, click chemistry, enzyme-mediated crosslinking, photopolymerization, and Schiff base reactions. These methods allow for precise control over the hydrogel's network structure and mechanical properties, which is particularly beneficial for tailoring hydrogels to meet the specific requirements of neural tissue engineering.
The Michael reaction is a mild cross-linking method that proceed under physiological and ner-neutral pH, making it well-suited for biological applications by selectively targeting biogenic amines [110]. This reaction involves a nucleophile adding to an unsaturated carbonyl compound, a process classified as conjugate addition, resulting in covalent bond formation [116]. PEG is commonly used to prepare injectable hydrogels through Michael addition, owing to its protein resistance, high hydrophilicity, and reduced susceptibility to degradation by mammalian enzymes [117]. However, recent studies have identified certain challenges with the Michael reaction, including a rapid polymerization rate that exceeds the manual mixing rate of components, resulting in non-uniform hydrogel networks and slow gelation rates [116,118,119].
Click chemistry is not a single, specific reaction; rather, it is a method for generating products that mimics natural examples [120]. These reactions are characterized by their minimal reactivity with cellular components and exceptional selectivity for linking molecular entities. Additionally, it is characterized by versatility, high yield, and rapid polymerization kinetics, all without the need for chemical additives, cytotoxic crosslinkers, or the generation of harmful byproducts [121]. This feature make click chemistry an attractive approach for the in situ fabrication of injectable hydrogels, enabling precise network formation under mild conditions [122,123]. This methodology encompasses diverse methods, including copper-catalyzed azide-alkyne cycloaddition and copper-free click reactions, such as Diels-Alder reactions, strain-promoted azide-alkyne cycloaddition, radical-mediated thiol-ene reactions, and oxime formation [120].
Enzyme-mediated crosslinking reactions are categorized as mild because most enzymes catalyze reaction at neutral pH and moderate temperatures. Thei high substrate specificity of enzymes minimizes undesired side reactions and the production of toxic substances [124,125]. A variety of enzymes are available to mediate crosslinking reactions for hydrogels synthesis, including transglutaminase, tyrosinase, ATP transferase, lysyl oxidase, plasma amine oxidase, phosphatases, thermolysin, beta-lactamase, and peroxidases [124]. Among these, horseradish peroxidase (HRP) stands out due to its excellent biocompatibility and rapid gelation capabilities, making it particularly effective for the in situ formation of injectable hydrogels [126].
Schiff base chemistry involves the reaction of aldehydes (or ketones) with amines to form imines (or Schiff bases), where the condensation of the carbonyl group and primary amine produces Schiff bases and water as a byproduct. This reversible reaction can occur under mild conditions [127] and is characterized by the formation of a rigid network with relatively slow gelation kinetics [128]. The pH sensitivity of Schiff bases further enables hydrogels to respond to biological stimuli dynamically. Moreover, by selecting different types of Schiff bases, researchers can fine-tune the mechanical properties and chemical stability of the hydrogels [129,130].
Photopolymerization involves the rapid and controllable conversion of liquid monomers or macromolecules into hydrogels via free radical polymerization under specific conditions. This method allows for the photopolymerization of hydrogels in vivo or in vitro using visible or ultraviolet light conjunction with a photoinitiator [[131], [132], [133]]. Photocrosslinked hydrogels typically demonstrate shorter gelation times, enhanced chemical stability, and improved mechanical strength compared to other crosslinking methods. However, the ultraviolet light exposure can induce cell death, and the exothermic nature of the reaction may damage surrounding cells, potentially leanding to localized tissue necrosis [131,[134], [135], [136], [137]]. Despite these challenges, recent advances in controlling crosslinking time and spatial location under physiological conditions have shown promising photopolymerization applications in the development of injectable hydrogels [135].
Physical crosslinking is driven by non-covalent interactions and molecular entanglement, including ionic interactions, hydrogen bonding, hydrophobic interactions, and crystallization. This approach avoids the use of chemical crosslinkers, often resulting in hydrogels with reversible and stimuli-responsive properties.
Ionic interaction crosslinked hydrogel structures are primarily formed by electrostatic interactions between oppositely charged molecules [138]. A prime example is alginate, a natural anionic polymer derived from brown seaweed, which is widely used in bioengineering due to its excellent biocompatibility, low toxicity, affordability, and mild gelation conditions. The common method for preparing hydrogels from alginate solutions involves mixing the solution with ionic crosslinkers, including divalent cations (e.g., Ca^2+^, Ba^2+^, and Sr^2+^). Calcium chloride (CaCl2) is among the most frequently used ionic crosslinkers for alginate; however, its high solubility in aqueous solutions, CaCl2 often results in rapid and uncontrolled gelation. To mitigate this issue, researchers have developed methods such as using phosphate-buffered solutions or incorporating substances with lower solubility [139,140].
Hydrophobic interactions, also known as hydrophobic bonds, represent a typical form of physical non-covalent crosslinking, playing a crucial role in the self-healing behavior of soft materials [13]. Hydrophobic interactions are the second type of interaction, distinct from hydrogen bonds, that are essential for the formation of biological structures. Typically, gel-forming molecules designed for hydrophobic interactions crosslinking contain both hydrophobic and hydrophilic segments. In the physical crosslinking of hydrogels, hydrophobic interactions generally exhibit greater strength than both hydrogen bonds and van der Waals forces, thereby contributing to the overall stability and functionality of the hydrogel [141].
Supramolecular chemistry, defined as "chemistry beyond the molecule," encompasses hydrogen bonding, metal chelation, hydrophobic interactions, π-π interactions, and van der Waals forces. This approach is predicated on the design of directional, reversible, and tunable molecular recognition motifs that allow for the formation of well-defined, self-assembled structures [142,143]. Supramolecular hydrogels exhibit excellent biocompatibility, biodegradability, and a high density of cell adhesion sites. However, these hydrogels also present several drawbacks, including rapid degradation, suboptimal mechanical compliance, and limited elasticity [144]. Recent advancements aimed at enhancing the mechanical strength of supramolecular hydrogels involve the development of double-network hydrogels, sliding-ring hydrogels, and microsphere composite hydrogels, all designed to enhance the mechanical strength and durability of supramolecular systems [145].
Hydrogen bond crosslinking involves the formation of crosslinked networks through interactions between hydrogen atoms and electronegative atoms, such as oxygen or nitrogen. This process primarily facilitates the fabrication of hydrogels, especially supramolecular hydrogels enhanced by multiple hydrogen bonds that demonstrate excellent toughness, self-healing capabilities, and recyclability [146,147], making them generally more robust than those produced via other physical crosslinking methods [148].
Chemical and physical crosslinking methods each possess distinct advantages and disadvantages. Chemical crosslinking involves the formation of permanent, irreversible covalent bonds between polymer chains, yielding hydrogels with improved stability and mechanical properties under physiological conditions [128]. However, these in vivo chemical reactions may lead to potential toxicity and undesired side reactions. In contrast, physical crosslinking involves transient, reversible non-covalent bonds [77]. Hydrogels produced via physical crosslinking are highly sensitive to external factors like temperature, pH, and ionic concentration, which simplifies their preparation. Additionally, physical gelation avoid the use of, monomers, or catalysts, thereby reducing related biocompatibility concerns and cytotoxicity; however, The dynamic and reversible nature of physical crosslinking can hinder precise control over degradation rates and behaviors, often resulting in lower overall stability compared to chemically crosslinked hydrogels [149]. In summary, chemical crosslinking offers robust and stable hydrogel networks suitable for demanding physiological environments, while physical crosslinking provides dynamic, responsive, and highly biocompatible systems. Future research should focus on optimizing these crosslinking strategies, potentially through hybrid approaches, to develop hydrogels that best meet the specific requirements of neural tissue engineering applications.
Physical Crosslinking. Hydrogen bonding, Guangzhao Zhang et al. prepared a hydrogel based on poly(ethylene oxide)-poly(N-isopropylacrylamide) (PEO-PNIPAm) through hydrogen bonding crosslinking [148]. Ionic interactions, Kun Zhang et al. prepared Hyaluronic acid/Sodium alginate (HA/SA) scaffolds through the interactions of SA, CaCO3, and HA. During the preparation process, the introduction of gluconic acid lactone (GDL), wherein the ratio of Ca2+ and COO was 1 : 2, and the molar ratio of Ca2+ and GDL was also 1 : 2. triggered a crosslinking reaction, resulting in HA/SA hydrogels after the reaction was complete [158]. Hydrophobic interactions, Hydrophobic interaction has been used to synthesize HA/methylcellulose (MC) hydrogels by M Douglas Baumann et al. [159]. Supramolecular chemistry, Biao Yang et al. used an appropriate combination of polypyrrole (PPy), gelatin (Gel), and agarose (Aga), and ultimately prepared supramolecular hydrogels through supramolecular chemical crosslinking [160].
Injectable hydrogel strategies for neural tissue engineering can be categorized into four endogenous repair, exogenous cell delivery or drug delivery, and combined delivery of exogenous cells with drugs agents [161].
Endogenous repair entails directly injecting functional hydrogels into the injury site to stimulate endogenous repair mechanisms, guiding tissue repair and regeneration, including nerve and vascular regeneration. Exogenous cell delivery involves encapsulating exogenous cells within injectable hydrogels, which provide a favorable environment upon injection into the injury site to facilitate repair. This process stimulates endogenous repair by releasing bioactive substances from the exogenous cells, achieving a synergistic effect in nerve tissue repair. Drug agent delivery utilizes injectable hydrogels as carriers to transport drugs directly to the injury site, enabling therapeutic effects through sustained or controlled release. Finally, combined delivery of cells and drugs agent allows injectable hydrogels to serve as dual carriers, enhancing repair and regeneration outcomes through synergistic effects. Essentially, injectable hydrogels actively participate in neural repair through biochemical, biomechanical, and immunomodulatory mechanisms. They can modulate the post-injury inflammatory microenvironment by promoting macrophage polarization toward the anti-inflammatory M2 phenotype and reducing oxidative stress, thereby creating a pro-regenerative milieu conducive to axonal growth and remyelination [58]. The biomimetic extracellular matrix structure and tunable stiffness of hydrogels regulate neural stem cell adhesion, migration, and lineage differentiation through relevant signaling pathways, promoting neuronal differentiation and synapse formation [162]. In addition, some aligned or conductive hydrogels provide directional and electrical cues that facilitate axonal extension and enhance functional connectivity [163]. Hydrogels that release growth factors or ions promote vascular and other functional reconstruction via specific signaling pathways, restoring neurovascular coupling and improving functional recovery [164]. These biologically and material-mediated synergistic mechanisms collectively underpin the therapeutic potential of injectable hydrogels in neural tissue engineering.
Table 1, Table 2, Table 3 summarize the progress of injectable hydrogels in neural tissue engineering and provides a detailed classification based on their applications in various nervous system injuries and diseases.Table 1Injectable hydrogels for brain tissue engineering in the central nervous system.Table 1Name/YearDieaseHydrogelAnimal profileIn vivo modelLoadedWithInjectionResultsCompositionPreparationTypeweightAgeAnestheticTypeSiteNeedle gradeTimeXiaoyin Liu 2023TBIDHCSelf-assembledmale SD rats200–220 gNonePentobarbital sodium (50 mg/kg)TBIBMEInto the TBI cavityNoneImmediately injectedThe DHC hydrogel significantly enhanced functional recovery in TBI rats by reducing glial scar formation, promoting structural remodeling, supporting NSC differentiation and recruitment, and facilitating angiogenesis via BME delivery.Yan Hu 2022Brain lesionsHA-PBA/Gel-Dopa hydrogelPH-responsiveC57BL/6J male miceNoneNone2 % Avertin (10 mg/mL)TBINoneInto the lesion22 G needleImmediately injectedAfter 3 weeks, it reduced glial scar formation, closed lesions, and enhanced neural cell infiltration.Ivana Perkucin 2022TBIAlginate/PEDOT hydrogelIon-sensitive crosslinking, Electrostatic interaction,Hydrophobic interactionC57BL/6Mouse22–35 g7–11weeks5 % isofluraneTBINoneInto the lesion22G needleImmediately injectedAfter 3 days, the hydrogel reduced neuroinflammation, matched brain tissue properties, provided effective electrical stimulation, and exhibited no cytotoxicity in 7-day neural cell tests.Ye Li 2022TBIFPGEGaSchiff base reaction, Electrostatic interaction, ThermosensitiveMale Wistar rats220–240 gNone10 % chloral hydrate, intraperitoneallyTBISHED-Exo (100 μg Exo into100 μL hydrogel)Into the lesionNoneNoneAfter 2 days, it reduced intracellular ROS formation; after 3 weeks, it enhanced neuroregeneration and restored motor functions.Y Wang 2023TBIFmoc-DDIKVAV hydrogelSelf-assembledMouse/C57BL/6NoneNone2 % isofluraneTBIMyoglobin (1 mg/mL) and/or cortical neural stem cells (5 × 104/μl cells)Into the striatumNoneImmediately injectedAfter 4 weeks, it supported stem cell engraftment, promoted vascularization, enhanced neuronal differentiation, and improved functional integration of grafts.Yantao Zheng 2021TBIGelMA-imidlight -responsive, Click chemistry cross-linkedRat (Sprague-Dawley)None8–11weeksIntraperitoneal, sodium pentobarbital (30 mg/kg)Cryogenic injuryPDA/SDF1α nanoparticles and hAMSCsInto the center of the damaged area26 G needleImmediately injectedAfter 2 weeks, it enhanced hAMSCs homing and neural differentiation, promoted endogenous nerve cell growth, and showed significant potential for TBI recovery.Jinrui Li 2021TBIGelatinEnzymatically cross-linkedMouse/C57BL/621–23 g6–8weeksIntraperitoneal, 10 % chloral hydrate (350 mg/kg)Moderate TBIBMSC (5 × 104 cells/mlInto the lesionNoneSeven days after TBI establishment,After 5 weeks, it improved neurogenesis and functional recovery in TBI mice, significantly increasing neural differentiation, cell viability, and neurotrophin secretion.Feng Qian 2021TBITM/PCSelf-assembledMouseNoneNoneIsoflurane (3.5 % induction and 1.5 % maintenanc)TBICurcuminUnder the endocraniumNoneImmediately injectedAfter 3 weeks, it decreased reactive astrocytes and activated microglia, exerted strong anti-inflammatory effects, and promoted nerve regeneration.Ru-Siou Hsu 2022TBIGelMAlight-assembledC57BL/6 miceNone7 weeks3 % isofluraneTBIGYBs /MBsInto the lesionNoneImmediately injectedThe conductive hydrogel promotes cell infiltration, anti-inflammation, and angiogenesis, while high-frequency magnetic fields enhance neurite growth and BDNF expression, boosting vascularization and neuron survival to restore brain connectivity.Luyu Wang 2022TBIhyaluronic acidEnzymatically cross-linkedC57BL/6 male mice22–25 gNoneNoneTBIBMSC and NGFInto the lesionNoneSeven days after TBI model establishmentThe 0.5 % HT hydrogel, exhibiting the lowest swelling ratio, was the optimal scaffold; in situ delivery of NGF- and BMSC-loaded hydrogel promoted motor and cognitive recovery and brain repair by supporting cell survival while reducing neuroinflammation and apoptosis.Name/YearDieaseHydrogelAnimal profileIn vivo modelLoadedWithInjectionResultsCompositionPreparationTypeweightAgeAnestheticTypeSiteNeedle gradeTimeYanhong Pei 2023Ischemic strokeChitosan-DF-PEG hydrogel/GelMA-PCL nanofiberSchiff base reactionRats/Sprague Dawley300–320 g7–8 weeks4.1 % pentobarbital sodium (40 mg/kg)Ischemic brain injury, middle cerebral artery occlusion (MCAO) modelBMSCsInto the perilesional site26 s-gauge, Hamilton, for 10 min24 h after MCAOAfter 2 weeks, significant neurogenesis and angiogenesis occurred, accompanied by reduced ischemic damage, infarct volume, microglial and astrocyte overactivation, and neurological deficits.Yuya Ohno 2023Ischemic strokeNcad-mRADASelf-assembledNSE-DTA miceNone8–16-weeks1 %–3 % isoflurane in oxygenIschemic stroke, middle cerebral artery occlusion (MCAO) modelNoneInto the siteGlass capillary needle, slowly stereotaxicallyinjected13 ± 1 day-post MCAO micce and 2 days-post-cryoinjured miceNeuroblast migration to the injured striatum promoted neuronal regeneration and functional recovery with cortical damage.Anup Tuladhar 2020Ischemic strokeHAMCPhysically cross-linkedRat/Sprague Dawley300–350 gNoneIsofluraneStroke (endothelin-1 model of focal ischemia)Cyclosporine and erythropoietinOnto the brain's cortical surfaceNone4 days after the stroke surgeryAfter 6 weeks, cyclosporine enhanced striatal plasticity, and erythropoietin stimulated endogenous NSPCs.Yang Liu 2023Ischemic strokeHG/dECMchemistry crosslinkedMale C57 mice25–30 gNoneIsofluraneCerebral focal ischemiaiPSC-NSCsinto the infarct cavityNone5 days after cerebral ischemiadECM composites support neural cell survival and differentiation, while HG hyaluronic acid hydrogels adjust to growth-related volume changes.Yi Liu 2020Intracerebral hemorrhageChitosan-HA (CH) hydrogelsSelf-assembledZebrafish & rat250–300 gNone200 ppm tricaine & isoflurane gaseous (5 % induction, 2 % maintenance)the zebrafish traumatic brain injury (TBI) and rat intracerebral hemorrhage (ICH) models.NoneInto the lesion (zebrafish) & into the striatum (rat)23-gauge needleImmediately injectedAfter 1 week, it promoted CNS healing and functional recovery by creating an adaptable microenvironment that supports NSC migration, proliferation, differentiation, and axonal growth within the brain injury cavity.Teck Chuan Lim 2020ICHGelatinEnzymatically cross-linkedSprague-Dawley rats250–300 gNoneNoneIntracerebral hemorrhage (via collagenase injection into the striatum)EGF (8 μg)Brain cavityNone2-week post-ICHAfter 4 weeks, sustained EGF release enhanced cell migration and promoted tissue regeneration.Jiaxin Zhang 2022ICHSFNoneCD1 mice20–25 gNoneisofluraneIntracerebral hemorrhageHydrogen sulfide (H2S)Into the siteNoneImmediately injectedH2S@SF hydrogel mitigated edema, hemorrhage, and pyroptosis, enhancing neurological recovery after ICH through sustained H2S release.Name/YearDieaseHydrogelAnimal profileIn vivo modelLoadedInjectionResultsWithCompositionPreparationTypeweightAgeAnestheticTypeSiteNeedle gradeTimeJuan Li 2016Parkinson's diseasePNIPAM based hydrogelMichael additionMale C57BL/6J miceNone8e10 weeks10 % chloralic hydrasParkinson's disease ratsactivin Binto the striatum of a mouse model of PD22-gauge needleImmediately injectedThe delivery system sustained activin B release for 5 weeks, enhancing neuroprotection, behavior, and biocompatibility in a PD mouse model.Helena Ferreira 2021multiple sclerosisHA-based hydrogelsNoneold male Wistar Han ratsNoneeight weeksA mixture of ketamine (75 mg⋅kg−1) and medetomidine (0.5 mg⋅kg−1)multiple sclerosisBMSCsthe lesion site25-gauge needleImmediately injectedThe BMSC-loaded hydrogel targets the corpus callosum, alleviating neuronal deficits, lowering clinical scores, and preventing disease relapse.Song Yi Lee 2020Alzheimer's Diseasehydrogels are based on HA-dpNonemaleNone5-weeks-oldNoneAlzheimer's Diseasedonepezilthe subcutaneous injectionsNoneNoneThe HA-dp/PD MS/FeSO4 hydrogel exhibited strong viscoelasticity, sustained release, and self-healing, lasting three weeks after injection, with improved pharmacokinetics over DPZ, providing a safe and effective Alzheimer's therapy.TBI: Traumatic brain injury; DHC: hyaluronan-collagen hydrogel; BME: bone marrow mesenchymal stem cell-derived exosomes; HA-PBA: phenylboronic acid grafted hyaluronic acid; Gel-Dopa: dopamine grafted gelatin; PEDOT: polymer, poly(3,4-ethylenedioxythiophene); FPGEGa: poly (cit-rate-gallic acid)-based multifunctional; Fmoc-DDIKVAV: the functional epitope encoding the binding domain of laminin; GelMA-imid: imidazole groups-modified gelatin methacrylate; TM/PC: post-trauma microenvironment-responsive, ROS depletion hydrogel embedded curcumin; GelMA:gelatin methacrylamide; GYBs:the electromagnetized gold nanoyarn balls; GYBs: magnetized gold nanoyarn balls; MBs: magnetic beads; BMSCs: bone mesenchymal stem cells; NGF: nerve growth factor.GC: glycol chitosan, DF-PEG: dibenzaldehyde-terminated polyethylene glycol, BMSCs: bone marrow mesenchymal stem cells, HAMC: hyaluronan and methylcellulose, HG: host–guest, dECM: decellularized extracellular matrix, iPSC-NSCs: induced pluripotent stem cell-derived neural stem cells, ICH: intracerebral hemorrhage, SF: silk fibroin.NSCs:Neural stem cells, PNIPAM: Poly(N-isopropylacrylamide), BMSCs: bone marrow mesenchymal stem cells, HA-dp: hyaluronic acid–dopamine.Table 2Injectable hydrogels for spinal cord tissue engineering in the central nervous system.Table 2Name/YearDieaseHydrogelAnimal profileIn vivo modelLoadedWithInjectionResultsCompositionPreparationTypeweightAgeAnestheticTypeSiteNeedle gradeTimeShangzhi Li 2022SCIAHA/DTPSchiff base cross-linkedRatNoneNoneNonespinal cord injuryNSCsInto the lesionNoneNoneAfter 1 week, it promoted NSC differentiation, bridged lesions, enhanced angiogenesis and remyelination, and improved neural regeneration and motor recovery.Yian Luo 2021SCIBOCPGSchiff base cross-linked and non-covalent crosslinkingAdult female SD rats280–300 gNonea mixture of xylazineand ketaminespinal cord injuryNSCsInto the lesionNoneAfter a weekEnhanced neuronal differentiation and axon outgrowth while suppressing astrocyte differentiation; after 6 weeks, it activated endogenous NSC neurogenesis and myelinated axon regeneration, promoting locomotor recovery.Luzhong Zhang 2022SCIIbuprofen-KYIGSRK conjugated hydrogelsMichael addition reactionfemale SD rats200 gNoneNoneSemi-transectionNoneInto the lesionNoneImmediately injectedEnhanced DRG neuron adhesion and neurite outgrowth; after 4 weeks, reduced glial scars, increased neurites, suppressed inflammation, and promoted motor recovery and nerve regeneration.Zengjie Zhang 2022SCIPLELThermosensitivefemale rats200 gNonePentobarbital solutionClip compressionEVsInto the lesionNoneImmediately injectedAfter 4 weeks, it reduced local inflammation and neural cell apoptosis, provided topological guidance for axon alignment, and promoted motor functional recovery.Zhiping Qi 2021SCIA dual-drug consisting injectable hydrogelThermosensitiveRat/Sprague-Dawley200–250 gNonePentobarbital solution (50 mg/kg)spinal cord injuryNSCsInto the lesionNoneImmediately injectedEnhanced neuronal differentiation and suppressed astrocyte differentiation; after 10 weeks, it promoted NSC integration, axonal regeneration, limited glial scar formation, and improved functional recovery.Gong Ho Han 2020SCITauroursodeoxycholic acidSchiff base reactionRat/Sprague-Dawley210–240gNoneIntraperitoneal injection,10 mg/kg Rompun & 50 mg/kg of Zoletil 50spinal cord injuryNoneInto the lesion26-gauge needleImmediately injectedAfter 1 week, it inhibited neuroinflammation.Kristyna Zaviskova 2018SCIHA-PH-RGD/FEnzymatically cross-linkedRat/Wistar250–300gNonePentobarbital anesthesia (60 mg/kg)HemisectionhWJ-MSCsInto the lesionNoneImmediately injectedAfter 8 weeks, it bridged the lesion cavity, promoted vascularization, and enhanced axonal sprouting, but did not improve locomotor recovery or blood vessel ingrowth, nor reduce glial scar density around the lesion.Jacob Matthews 2021SCIBC/FB gelThermosensitiveRat/Sprague-Dawley170–220 gNoneBuprenorphine, 5 % Isofluo rane, maintenance 2 % isofluoraneThe dorsal column was crushed bilaterallyDecorinInto the lesionNoneImmediately injectedAfter 6 weeks, cavitation was prevented, lesion sites were remodeled, regeneration genes upregulated, axon growth increased, and electrophysiological, locomotor, and sensory functions improved.Jiyun Cheng 2021SCIGelMAUV-sensitive crosslinkingSprague–Dawley (SD) ratsNoneNoneNonespinal cord injuryBone MSCs-derived exosomesInto the lesionNoneImmediately injectedAfter 6 weeks, it promoted neurological functional recovery; after 7 days of cell culture, it induced Tuj-1-positive neuron differentiation, reduced astrocyte scars, and increased axonal elongation.Kiet A Tran 2022SCIRADA-16ISelf-assembledRat/Sprague-Dawley225–250 gNoneIsoflurane 5 %ContusionhMSCsInto the lesionNoneTwo weeks after injuryAfter 2 weeks, it promoted axon elongation at the spinal cord lesion and provided topological cues guiding axon alignment.Laura M Marquardt 2020SCISHIELD:Michael addition reaction; Thermosensitive; Self-assembledRat/Fischer 344 (F344)NoneNone1–3 % isoflurane in oxygenContusionSchwann cell (4.5 × 105 cells)Into the lesion33-gauge Hamilton syringeTwo weeks after injuryAfter 4 weeks, it reduced neuronal damage, secondary injury, and cystic cavitation, enhanced cell transplantation outcomes, and improved function through increased forelimb strength and coordination.Shiyu Chen 2020SCISF/PDASchiff base reaction;Self-polymerizationRat/Sprague Dawley220 gage of 17–18 days10 % chloral hydrate solution (0.3 mL/100 g)3 mm lateral hemisectionNoneInto the injury siteNoneNoneAfter 2 weeks, it enhanced neuronal development and promoted healing in spinal cord injury.R Chase Cornelison 2018SCIAcellular peripheral nerve graftThermosensitiveRat/Sprague Dawley250–300 gNoneIsofluraneContusionNoneInto the injury siteUltraMicroPump3 microinjection systemOne week after injuryAfter 1 week, it modulated macrophage polarization; after 8 weeks, it enhanced neuronal regeneration and axon coverage as an effective, minimally invasive SCI therapy substrate.Kaijia Zhang 2020SCIPD/GONon-covalent crosslinkingFemale Sprague-Dawley rats190–220 gNoneIsofluraneModerate crushingDiacereinIn situNoneImmediately injectedAfter 4 weeks, it regulated astrocyte hyperactivity and inflammation in vitro and promoted SCI repair in vivo.Christy Kwokdinata 2023SCIHyaluronan-gelatin hydrogelsNoneFemale Sprague–Dawley rats230–250 g8–9Wketamine (50 mg/kg) and xylazine (5 mg/kg)A dorsal laminectomyspinal cord progenitor cells (SCPCs)In situ701 N Hamilton syringeImmediately injectedHydrogel encapsulation enhanced SCPC survival and retention after spinal cord implantation, promoting differentiation into neuronal and oligodendroglial lineages.Heng Zhou 2023SCIGelatin methacryloyl hydrogelNoneFemale Sprague-Dawley ratsNone4 weeks old3 % pheno-barbitalspinal cord injuryZIF-8In situstereotaxic instrument and microinjection pump24 h post injuryZIF-8 activates JNK1 and p38 MAPK to enhance DPSC neurodifferentiation and angiogenesis, surpassing Zn^2+^ alone; when implanted in SCI, it restores Zn levels, elevates VEGF-A secretion, and supports vascular and neural regeneration.Farzaneh Sorouri 2023SCIHSP-F/BCSMichael-type additionAdult male Wistar rats250–280 gNonea ketamine (100 mg/kg)–xylazine (10 mg/kg) cocktailspinal cord injuryNoneIn situ27-gauge Hamilton syringeOne week after injuryThe HSP-F/BCS hydrogel offers dual delivery, protecting BCS nanoparticles from rapid degradation, preventing burst release of Buc, and promoting localized delivery of FA and Buc.Heng Wang 2023SCIGelMAPhoto/light-responsiveSprague-Dawley (SD) female rats220 g8-week-oldsodium pentobarbital administered intraperitoneallyspinal cord injuryBerberine (Ber)In situNoneOne week after injuryGelMA hydrogel enables minimally invasive delivery and sustained sEVs-Ber release, reducing inflammation and fibrosis to promote nerve regeneration and motor recovery after SCI.Tiemei Liu 2023SCISAO and NH2-Gel-PANINoneFemale Sprague Dawley (SD) ratsNoneNoneNonespinal cord injuryNSCs and DPLIn situNoneImmediately injectedDPL-loaded hydrogels enhanced NSC proliferation and differentiation, promoting neuronal integration and reducing glial scars in SCI rats.Fang Wang 2023SCIPLL hydrogelsPhysical cross-linkingRatsNoneNonePentobarbital sodiumspinal cord injuryLiClIn situNoneImmediately injectedLiCl-PLL hydrogel promoted NSC neuronal differentiation, reduced oxidative stress and inflammation, inhibited glial scars, enhanced axon growth and synapse repair, accelerating SCI recovery.Dun Liu 2023SCICeNP-GelNoneFemaleSD rats200–220 gNoneisoflurane inhalation anesthesiaspinal cord injuryNSCsthe lesion siteNoneImmediately injectedCeNP-Gel alleviated oxidative stress and induced M2 microglia polarization to reduce inflammation, thereby enhancing NSC viability.Hou Liu 2023SCIICH/NSCsNoneFemale Sprague−Dawley rats250−350 gNonepentobarbitalsodiumspinal cord injuryNSCsthe lesion siteNoneImmediately injectedNSC-loaded hydrogel scaffold promoted axon regeneration, reduced cavities and glial scars, and enhanced locomotor recovery within 6 weeks.Shangzhi Li 2022SCIAHA/DTPSchiff base reactionRats220–250 gNone50 mg/kg sodium pentobarbitalspinal cord injuryNonethe lesion siteNoneImmediately injectedThe hydrogels exhibited self-healing, injectability, and biocompatibility, promoting NSC differentiation, tissue regeneration, and locomotor recovery in SCI rats.Linquan Zhou 2022SCIMethacrylate-silk fibroin hydrogelsUV photocrosslinkingRatsNoneNoneintraperitoneal injectionof pentobarbital anesthesiaspinal cord injurybFGFthe lesion siteNoneImmediately injectedSustained bFGF release from the hydrogel improved mitochondrial function, inhibited glial proliferation, and promoted axon regeneration in vitro and in vivo.Kiet A. Tran 2022SCIMagneti-cally-responsive hydrogelsSelf-assemblingFemaleSprague-Dawley rats225–250 gNone5 % of isofluranespinal cord injuryNPCthe lesion siteNoneTwo weeks after injuryIn vitro and in vivo, MP-seeded scaffolds supported hMSC and NPC survival, promoted axon growth and host fiber infiltration, and magnetic alignment post-injection rapidly improved fiber orientation.He Z 2018SCIHAMC hydrogelsNoneAdult female SpragueDawley rats200–230 gNoneanesthetized with a nitrous oxide/oxygen mixture (70 %/30 %) containing 1.4 % isofluranespinal cord injuryNonethe lesion siteNoneImmediately injectedThe HAMC-KAFAK/BDNF hydrogel promotes functional recovery in SCI rats by modulating inflammation and enhancing axonal regeneration.XiaoLing Li 2019SCICab-M/HNoneAdult female SpragueDawley rats200–220 gNonechloral hydrate solu-tion (10 wt%, 0.3 mL/100 g)spinal cord injuryCabthe lesion siteNoneImmediately injectedIn a rat lateral hemisection SCI model, a single local Cab-M/H application significantly reduced inhibitory scarring, enhanced axonal growth, and improved sensorimotor function.Chong Wang 2018SCIGCP-hydrogelUV photocrosslinkingFemale C57 mice20 g10-week-oldpentasorbital sodium (0.1 %)spinal cord injuryESCsthe lesion siteNoneImmediately injectedESC-loaded hydrogels injected into SCI mice enhance tissue repair, angiogenesis, and neural marker expression (Tuj1/MAP2/Syn) while reducing inflammation.SCI: spinal cord injury, AHA/DTP: Aldehyde-functionalized Hyaluronic Acid /Dithiopropionic acid-based hydrogel, NSCs: neural stem cells, BOCPG: BOC-doped polypyrrole with gelatin hydrogels, PLEL: poly(d,l-lactide)-poly(ethylene glycol)-poly(d,l-lactide), EVs: extracellular vesicles, HA-PH: hydroxyphenyl derivative of hyaluronic acid, hWJ-MSCs: human Wharton's jelly-derived mesenchymal stem cells, BC/FB: bovine collagen/fibrinogen, hMSCs: human mesenchymal stem cells, SF/PDA: silk fibroin/polydopamine, PD: diacerein (4arm-PEG-diacerein), GO: graphene oxide, ZIF-8: The zeolitic imidazolate framework 8, HASH-FA: hyaluronic acid modified with ferulic acid, BCS NPs: Bucladesine-encapsulated chitosan nanoparticles, SAO: sodium hyaluronate oxide, NH2-Gel-PANI: polyaniline-grafted gelatin, DPL: donepezil, PLL poly(lipoic acid-co-sodium lipoatehydrogels, LiCl: Lithium chloride, CeO2@BSA nanoparticles (CeNPs): albumin biomimetic cerium oxide nanoparticles, AHA: aldehyde-modified hyaluronic acid, DTP: 3,3′-dithiobis(propionylhydrazide), bFGF: Basic fibroblast growth factor, NPC: neural progenitor cells, HAMC: Hyaluronan-methylcellulose, Cab-M/H: Cab-M encapsulated hybrid hydrogel, Cab: Cabazitaxel, β-CD: beta-cyclodextrin, GCP-hydroge: Gelatin-acrylated β-cyclodextrin polyethylene glycol hydrogel, ESCs: embryonic stem cells.Table 3Injectable hydrogels in the peripheral nervous system.Table 3Name/YearDieaseHydrogelAnimal profileIn vivo modelLoadedWithInjectionResultsCompositionPreparationTypeweightAgeAnestheticTypeSiteNeedle gradeTimeSaba Nemati Mahand 2023Sciatic nerve injuryPEtOx hydrogelphoto-crosslinkedWistar rats250–270 g3 months oldKetamine 5 % and xylazine 2 %Injure the sciatic nerve4-Aminopyridine (4-AP) small moleculesThe lesion siteNoneImmediately injectedThe hydrogel exhibited excellent biocompatibility, swelling, degradation, and hemocompatibility. In vitro, GelMA/PEtOx enhanced cell proliferation versus control, with GelMA/PEtOx + 4-AP performing best.Zhenwei Yi 2023Sciatic nerve injuryALHA/CMCS/CP (ACCP) hydrogelDynamic Schiff base bonding and electrostatic interactionsMale Sprague-Dawley (SD) ratsAverage weight 225 gNone3 % isofluraneSciatic nerve crush injuryNoneThe lesion site25 GneedleImmediately injectedThe Schiff base–derived ACCP3 hydrogel matches the sciatic nerve's mechanical and electrical properties and, when injected into crush injuries, enhances nerve conduction, promotes axonal regeneration and myelination, and prevents muscle atrophy.Hongyun Xuan 2023Sciatic nerve injuryHASPy hydrogelNoneFemale Sprague Dawley (SD) rats200–250 g8 weeksChloral hydrate sodium pentobarbital mixtureInjure the sciatic nerveNoneThe lesion site1 mL syringeImmediately injectedDirect injection of the hydrogel into rat sciatic nerve crush injuries promotes in vivo nerve regeneration, functional recovery, and remyelination.Xiaodie Ma 2023Diabetic peripheral neuropathy (DPN)Thermoresponsive Hybrid HydrogelGO–polymer hybrid hydrogelNoneNoneNoneNoneNoneThe synergistic effect of exosomesThe lesion siteNoneNoneCell experiments confirmed the GO–thermoresponsive polymer hydrogel's biocompatibility, while exosome loading enabled NIR-triggered release and synergistically promoted Schwann cell proliferation and migration.Yunfan Kong 2023NPboronic ester-basedhydrogelNonemale C57BL/6 miceNine-week-oldNoneIsofluraneInjure the sciatic nerveCurcuminThe lesion siteNoneImmediately injectedThis study suggests orthotopic Gel-Cur-M injection as a promising therapy, particularly for peripheral neuropathy patients needing surgery.Zahra Taherian Mobarakeh 2023Sciatic nerve injuryFibrin hydrogelNoneMale Wistar rats240–280g4 months oldKetamine + XylazineInjure the sciatic nervehEnSCsThe lesion site18-gage needleImmediately injectedThe hydrogel scaffolds with insulin nanoparticles and hEnSCs show potential as biomaterials for sciatic nerve regeneration.Wanlin Xu 2023Sciatic nerve injuryQK-NLs@Gel hydrogelUV photocrosslinkingSprague-Dawley rats200–220g6–8 weeksNoneInjure the sciatic nerveVEGFThe lesion siteNoneImmediately injectedFour weeks post-surgery, rats injected with QK-NLs@Gel showed enhanced axonal regeneration, remyelination, and superior functional recovery compared to other groups in vivo.Yingfan Chen 2023Sciatic nerve injuryR-phage/FG hydrogelIonic interactionsMale SD rats280–300 gNoneisofluraneInjure the sciatic nerveNoneThe lesion siteNoneImmediately injectedThe R-phage/FG fiber shows promise as an alternative to autologous nerve grafts for large nerve injury repair, offering a novel phage-based biomaterial that promotes neurogenesis in PNI treatment.Vahid Zolfagharzadeh 2023Sciatic nerve injuryHAPh/ColPh microspheresChemical crosslinkingAdult male Wistar ratsNoneNoneKetamine + XylazineInjure the sciatic nerveInsulinThe lesion site22-gauge needleImmediately injectedInsulin-loaded biomimetic microspheres effectively enhanced nerve regeneration in rats.Shuai Liu 2023CNI-EDHGNoneMale Sprague-Dawley rats12 weeks old12 weeks oldNoneRat with CNI-EDADSC-ExoThe lesion siteNoneImmediately injectedApplication of HG@Exo at the nerve injury site significantly improves erectile function in a bilateral cavernous nerve injury rat model.Wanlin Xu 2021Sciatic nerve injuryMethacrylic anhydride hydrogelUV photocrosslinkingMale specific-pathogen-free Sprague-Dawley rats6–8 weeks200–220 g2 % pentobarbital sodiumRat sciatic nerve crush injury modelVascular endothelial growth factor (VEGF)The lesion siteNoneImmediately injectedThe modified gelatin provided sustained VEGF delivery and accelerated crushed peripheral nerve repair.Zhi Yao 2022Sciatic nerve injuryBisphosphonate-based injectable hydrogelChemical crosslinkingFemale SD rats2 months200–250gKetamine + xylazineInjure the sciatic nerveMgThe lesion siteNoneImmediately injectedThe Mg^2+^-releasing hydrogel combined with a 3D-engineered PCL conduit promotes peripheral nerve regeneration and functional recovery, offering a novel strategy for repairing challenging nerve injuries.Travis A. Prest 2021Sciatic nerve injuryNerve-tissue-specific hydrogelNoneLewis ratsNoneNoneAnesthetizedInjure the sciatic nerveNoneThe lesion site25G needleImmediately injectedHydrogel injection post-nerve crush modestly improved function and accelerated recovery, with increased axon counts despite unchanged motor action potentials.Tania L. Lopez-Silva 2020Sciatic nerve injuryMultidomain peptide (MDP) hydrogelsSelf-assemblingFemale Sprague-Dawley ratsAverage weight 225 gNoneIsofluraneInjure the sciatic nerveNoneThe lesion siteNoneImmediately injectedMDPs promote neurite outgrowth and stimulate a multicellular regenerative response in peripheral nerve injury.Kangzhen Chen 2022Brachial plexus avulsion (BPA)Pluronic F-127 hydrogelNoneAdult female Sprague–Dawley (SD) rats180–220 g8–10 weeksPentobarbital sodiumBPA-reimplantation modelsNOGO receptor-interacting protein 1 (LINGO-1)The lesion site10 μL Hamilton syringeFirst day after surgeryHydrogel-delivered miR-615 agomir suppresses LINGO-1, promoting nerve regeneration, angiogenesis, and motor recovery while reducing neuronal loss and astrocyte activation.Yunfan Kong 2021Sciatic nerve injuryHA-PVA-Hep hydrogelNoneC57BL/6 mice18–20 g12 weeks oldIsofluraneInjure the sciatic nerveNoneThe lesion siteNoneImmediately injectedThe hydrogel improved sensorimotor function, prevented muscle atrophy, protected neurons, and promoted nerve regeneration and remyelination 28 days post-injury.Majid Rahmati 22021Sciatic nerve injuryAlg/Chit hydrogelIonic interactionsMale Wistar rats250–270gNoneKetamine 5 %/Xylazine 2 %Injure the sciatic nerveBerberineThe lesion siteNoneImmediately injectedBerberine, particularly at 1 % in hydrogel, enhances sciatic nerve regeneration, with Alg/Chit hydrogel demonstrating therapeutic potential for peripheral nerve injuries.Majid Salehi 2018Sciatic nerve injuryAlginate/chitosan hydrogelChemical crosslinkingAdult male Wistar rats250–270 g3 months oldKetamine 5 %/xylazine 2 %Injure the sciatic nerveOE‐MSCsThe lesion site18‐Gauge needleImmediately injectedAlg/Chit hydrogel with OE-MSCs enhances sciatic nerve regeneration compared to controls and cell-free hydrogel.PEtOx: poly(2-ethyl-2-oxazoline), ALHA: aldehyde-based hyaluronic acid, CMCS: carboxymethyl chitosan, PPy: polypyrrole, NP: neuropathic pain, hEnSCs: human endometrial stem cells, QK-NLs: QK-encapsulated nanoliposomes, VEGF: vascular endothelial growth factor, FG: fibrin gel, R-phage: RGD-displaying phage, ColPh: phenol-substituted collagen, HAPh: phenol-substituted hyaluronic acid, CNI-ED: cavernous nerve injury-induced erectile dysfunction, HG: hydroxyethyl chitosan/sodium β-glycerophosphate hydrogel, ADSC-Exo: adipose mesenchymal stem cell-derived exosomes, HA-PVA-Hep: hyaluronic acid-phenylboronic acid-poly(vinyl alcohol)-heparin, Alg/Chit: alginate/chitosan, OE-MSCs: olfactory ectomesenchymal stem cells.
TBI is one of the most prevalent forms of brain injury globally, with high incidence and mortality rates that pose significant clinical and public health challenges [[165], [166], [167]]. TBI typically results in outcomes such as hemiplegia, aphasia, and death, primarily due to neuronal loss, extensive vascular damage, and tissue necrosis [153,[168], [169], [170]].
The pathological procession of TBI comprises both primary and secondary injuries. Primary brain injury refers to immediate mechanical damage, inflicted at the moment of trauma, which can lead to contusions, lacerations, intracranial hemorrhage, and skull deformation [171]. In contrast, secondary brain injury, or delayed non-mechanical injury, is a continuous pathological process that starts at the time of injury but manifests later through delayed clinical symptoms. It encompasses inflammation, oxidative stress, excitotoxicity, and subsequent cell death [172]. Although primary brain injury is untreatable, secondary brain injury exacerbates damage to brain structure and function, making its prevention and mitigation the primary therapeutic goal [173,174].
In recent years, the growing attention to TBI has spurred extensive research and led to refined clinical interventions, including surgery, pharmacotherapy, and rehabilitation. However, these conventional treatments primarily alleviate symptoms rather than addressing the underlying tissue damage, thereby hindering the potential for complete functional recovery [175]. Surgical interventions remain complex and fraught with uncontrollable factors, potentially leading to severe complications like hydrocephalus and epilepsy [176]. Similarly, pharmacotherapy and rehabilitation face significant challenges due to the BBB, which restricts the effective delivery of many therapeutic agents to the injured areas, thus compromising treatment efficacy [177]. Therefore, exploring new therapeutic strategies is urgent.
Therefore, injectable hydrogels, as versatile carriers for cell and drug delivery, offer significant advantages as novel biomaterials. They exhibit excellent adaptability, biocompatibility, and biodegradability [16]. Through minimally invasive administration, injectable hydrogels can effectively fill injury cavities and modulate the host neural tissue by promoting cell adhesion, neurite extension, proliferation, and differentiation [178,179].
Yan Hu et al. [180] (Fig. 4.1) suggested that challenges in neural repair following TBI stem from mismatches in composition, stiffness, and viscoelasticity between hydrogels and brain tissue. These challenges require more rational hydrogel design strategies focused on optimizing mechanical properties (stiffness and viscoelasticity) to improve TBI treatment outcomes. In their study, the authors developed a self-healing injectable hydrogel composed of phenylboronic acid-grafted hyaluronic acid (HA-PBA) and dopamine-grafted gelatin (Gel-Dopa). When directly injected into the brain injury site of TBI rats, this hydrogel reduced glial scarring, sealed the lesion, and enhance neuronal infiltration within three weeks. These findings suggest that endogenous repair of TBI using hydrogels derived solely from these materials is a feasible therapeutic approach.Fig. 4Traumatic brain injury (TBI).Fig. 4
The introduction of stem cells into the brain injury site offers another promising strategy to promote neural repair. Once delivered, stem cells interact with local cells and biochemical signals to enhance the secretion of neurotrophic factors, inhibit neuroinflammation, promote synaptogenesis in injured neurons, and release neurotransmitters [181,182]. Among the various stem cell types, bone marrow mesenchymal stem cells (BMSCs) are particularly popular used for their therapeutic potential. Luyu Wang et al. [154] (Fig. 4.2) developed an injectable HT hydrogel through a dual-enzyme crosslinking method utilizing GalOx and HRP. This hydrogel, encapsulating both BMSCs and nerve growth factor (NGF), was administered to mice with TBI. The treatment enhanced the recovery of motor, learning, and memory functions, and accelerated the healing process of the damaged brain tissue.
Following TBI, the injured brain releases substantial amounts of reactive oxygen species (ROS) which accumulate and create a toxic environment for neurons, ultimately leading to neuronal death and additional functional deficits. Feng Qian et al. [183] investigated neuroprotective potential of curcumin (Cur), a compound recognized for its ROS-scavenging properties. They developed an injectable, ROS-scavenging hydrogel containing curcumin, referred to as TM/PC. This hydrogel can be directly injected into the injury site, not only effectively bypassing the BBB and facilitating localized drug accumulation, but also provides a sustained release of curcumin. In their mouse model of TBI, the injection of the TM/PC hydrogel enhanced tissue regeneration, inhibited tissue loss and liquefaction, and promoted overall functional recovery.
In parallel, emerging research highlights the promise of exosomes derived from MSCs in acellular therapies for various injuries [184,185]. These exosomes serve as nano-carriers for proteins, RNA, and lipids, have been shown to promote axonal regeneration, enhance angiogenesis, and reduce inflammation, benefits that are highly desirable for TBI repair [186]. Recently, Xiaoyin Liu et al. [187] (Fig. 4.3) sought to identify biomaterials capable of incorporating exosomes for controlled, sustained release. They employed a hyaluronic acid-collagen hydrogel loaded with BMSC-derived exosomes (BME), which was directly injected into the pre-modeled brain injury cavity in TBI rats. The combination of (hyaluronan-collagen hydrogel) DHC hydrogel and BME was found to inhibit glial scar formation, promote brain structural remodeling, enhance NSC recruitment, and facilitate neuronal regeneration. These effects contributed to significant improvements in spatial learning, memory, motor, and sensory functions following brain injury.
Yan Hu 2022 [180] (Endogenous Repair)a1)Schematic of preparing an endogenous repair hydrogel.b1)Injection of the hydrogel into TBI rat brains.c1)Astrocyte proliferation around the lesion cavity and neuronal migration into the cavity, illustrating endogenous tissue repair.
Luyu Wang 2022 [154] (Exogenous Cell Delivery)a2)Physical characterization of HT hydrogel (gelation time, water content, rheological modulus).b2)Functional outcomes measured by mNSS scores in TBI mice.c2)Representative images showing lesion volumes after 28 days, highlighting the efficacy of cell-loaded hydrogels.
Xiaoyin Liu 2023 [187] (Drug Agent Delivery)a3)Overall schematic of DHA hydrogel application.b3)Preparation process of DHA hydrogel.c3)Three-step crosslinking in DHC-BME.d3)DHA-BME promotes angiogenesis and neurogenesis, demonstrating its therapeutic potential for TBI repair.
Stroke is a leading cause of death and disability globally, posing considerable economic and social challenges for patients. Ischemic stroke is the most prevalent type, representing over 62 % of all strokes [188]. Ischemic stroke is primarily characterized by transient or permanent focal cerebral ischemia due to vascular occlusion [189,190].
From a pathophysiological standpoint, the ischemic cascade in ischemic stroke encompasses a series of neurochemical processes. This phenomenon accounts for why cells with varying degrees of ischemia may undergo distinct chemical processes. The cascade can be summarized as focal cerebral hypoperfusion results in cellular bioenergetic failure, followed by excitotoxicity, oxidative stress, blood-brain barrier dysfunction, and post-ischemic inflammation. These processes ultimately lead to the death of neurons, glial cells, and endothelial cells, resulting in impaired brain function and severe neurological and motor deficits [191,192].
As technology advances, treatment methods have diversified to include reperfusion therapy, pharmacotherapy, and rehabilitation therapy; however, outcomes frequently fall short of expectations. This is primarily because traditional treatments focus on alleviating symptoms, failing to promote neuroregeneration or recovery of neurological function [193]. Consequently, research on biomaterials for delivering therapeutic agents to damaged brain tissue has gained considerable momentum, with injectable hydrogels recognized as a leading strategy.
Yuya Ohno et al. [194] (Fig. 5.1) propose that using biomaterials to enhance the migration of young neurons (neuroblasts) derived from endogenous neural stem cells is a promising strategy. After a stroke, neuroblasts are attracted to chemokines and migrate to the lesion site, where they regenerate neurons, facilitating functional recovery [195,196]. Typically, neuroblasts use adjacent neuroblasts, blood vessels, and radial glial cells as scaffolds for migration. However, neuroblast migration can be inhibited by the hypertrophy of reactive astrocytes following brain injury. This is one reason why repairing stroke-induced brain damage is challenging. To address this, Yuya Ohno et al. [194] developed a the extracellular domain of N-cadherin labeled with mRADA (Ncad-mRADA), which can stably bind to mRADA hydrogels. They directly injected Ncad-mRADA hydrogels into the lesion sites of modeled ischemic stroke mice to enable migrating neuroblasts to contact the fibrous Ncad-mRADA hydrogels and effectively migrate to the injured striatum (a deep brain region). Ultimately, this approach was found to enhance neuronal regeneration and improve functional recovery following cortical brain injury.Fig. 5Ischemic stroke.Fig. 5
Stem cell therapy targeting the repair of infarcted tissue and enhancing functional recovery through nutritional support or direct replacement of damaged neurons has long been a focus in the treatment of ischemic stroke [197,198]. However, the efficacy of stem cell therapy for ischemic stroke is significantly compromised by insufficient migration of stem cells to the infarcted area, high mortality rates of transplanted cells, and poor integration of these cells into the damaged brain tissue [199,200]. In recent studies, Yang Liu et al. [201] (Fig. 5.2) developed a composite hydrogel composed of decellularized extracellular matrix (dECM) and matrix metalloproteinase (MMP)-responsive host-guest (HG) hydrogel (HG/dECM composite hydrogel), incorporating induced pluripotent stem cell-derived neural stem cells (iPSC-NSC). The reversible crosslinked dynamic network of the HG hydrogel not only reduces compression of the encapsulated NSCs during injection but also accommodates substantial volume expansion during cell proliferation. Additionally, the HG/dECM composite hydrogel provides various ECM proteins that sustain NSC viability and growth while establishing a favorable microenvironment for neuronal processes. Ultimately, they directly injected this hydrogel into the infarct cavity of a mouse model of focal cerebral ischemia, finding that it effectively addressed issues related to cell retention, survival, and poor differentiation of transplanted stem cells after ischemic stroke.
Cyclosporine A (CsA), a widely used immunosuppressant, has been shown to stimulate the generation of endogenous stem cells in the brains of animals. Erythropoietin (EPO) promotes erythrocyte production and has been demonstrated to enhance neurogenesis in rodent brains. Importantly, both drugs exhibit neuroprotective effects when administered within 24 h after injury [202]. Consequently, Nup Tuladhar et al. [202] (Fig. 5.3) developed a biocompatible hydrogel composed of hyaluronic acid and methylcellulose (HAMC) to deliver cyclosporine and erythropoietin directly to the cortical surface of a pre-established mouse model of stroke. The results indicated that each drug had distinct effects on brain CsA increased plasticity in the striatum, while EPO stimulated endogenous neural stem/progenitor cells (nspc). Interestingly, the combined administration of both drugs accelerated functional recovery and enhanced neural tissue repair.
Yuya Ohno 2023 [194] (Endogenous Repair)a1)Production of Ncad-mRADA hydrogel.b1)Injection into the ischemic site.c1)Labeling of V-SVZ-derived cells for histological and neurological assessment, illustrating endogenous repair responses.
Yang Liu 2023 [201] (Exogenous Cell Delivery)a2)Design of MMP-responsive host-guest (HG) hydrogel.b2)HG/dECM composite hydrogel schematic.c2)HG/dECM promotes survival and differentiation of iPSC-derived neural stem cells in vivo, supporting functional regeneration.
Nup Tuladhar 2020 [202] (Drug Agent Delivery)a3)Injection of HAMC hydrogel loaded with therapeutic agents onto damaged brain surface.b3)Sustained local release of drugs in ischemic tissue.c3)Combined CsA and EPO delivery via hydrogel reduces infarct volume, demonstrating therapeutic efficacy.
Intracerebral hemorrhage (ICH) accounts for approximately 10–15 % of all strokes and is the second most common subtype after ischemic stroke [203] It is also the deadliest form of hemorrhagic stroke due to its high mortality rate [204]. In roughly two-thirds of primary ICH cases, patients have either a history of hypertension or are newly diagnosed with the condition, with hypertensive small vessel disease serving as the predominant underlying mechanism. This vascular pathology can lead to the formation of small lipohyalinotic aneurysms that may eventually rupture, causing cerebral hemorrhage [205]. The brain injury associated with ICH can be divided into primary and secondary components. Primary injury results directly from the mass effect and physical disruption caused by the hematoma, while secondary injury develops as a consequence of deleterious factors released from the clot that further damage the surrounding brain tissue [206,207]. Although ICH is primarily a vascular disease, the rapid deterioration of neurological function can occur due to hematoma expansion or the rupture of the ventricular system.
Treatment strategies for ICH generally involve both surgical and conservative approaches. Surgical intervention aims to reduce hemorrhage volume and mitigate neuronal damage by alleviating local ischemia and removing toxic substances [208]. However, surgical treatments face significant challenges, including incomplete resection due to the inaccessibility of lesions, high rates of post-surgical rebleeding, and the potential for secondary brain injury induced by the procedure itself [209]. On the other hand, conservative management primarily relies on pharmacotherapy to stabilize patients during the perioperative period and limit further damage. Unfortunately, the blood-brain barrier restricts the effective delivery of most drugs to the injured brain regions, and extensive research has shown that these pharmacological treatments do not significantly improve clinical outcomes [210]. In this context, injectable hydrogels offer a promising alternative approach for developing neuroprotective therapies for ICH-related brain injury. Hydrogels can serve as efficient carriers for drug delivery and may provide targeted, sustained release of therapeutic agents, potentially overcoming some of the limitations of current treatment modalities.
Self-healing hydrogels are emerging as promising candidates for neural therapies because they can be directly injected into damaged regions using minimally invasive techniques, thereby effectively filling tissue defects once they have transitioned to a stable gel state. Yi Liu et al. [86] (Fig. 6.1) developed a self-healing injectable hydrogel by incorporating hyaluronic acid into a chitosan-based system to form a SIPN. They hypothesized that this HA-containing SIPN hydrogel would enhance biocompatibility and promote repair in CNS injuries. In a rat model of ICH, direct injection of the hydrogel into the striatum created a favorable microenvironment that supported the migration, proliferation, and differentiation of NSCs while promoting axonal development. These combined effects contributed to improved healing and functional recovery in the CNS.Fig. 6Hemorrhagic stroke (intracerebral hemorrhage).Fig. 6
Hydrogen sulfide (H2S) plays a crucial neuroprotective role as an endogenous signaling molecule in the CNS. Jiaxin Zhang et al. [211] (Fig. 6.2) addressed the challenge of sustained H2S delivery by developing a silk fibroin (SF) hydrogel loaded with H2S. Using stereotactic in situ injection in a severe ICH mouse model, the H2S@SF hydrogel provided a sustained release of H2S directly at the injury site, effectively overcoming the blood-brain barrier. This approach significantly reduced neuronal apoptosis in the striatum, cortex, and hippocampus, enhancing neurological recovery following cerebral hemorrhage.
In a complementary strategy, Teck Chuan Lim et al. [212] developed an injectable gelatin-based hydrogel that was enzymatically cross-linked and loaded with epidermal growth factor (EGF), referred to as Gtn-EGF. In their intracerebral hemorrhage model, the Gtn-EGF hydrogel was injected two weeks after the injury—when residual cavities were still present—and evaluated one month post-hemorrhage. The results demonstrated that the hydrogel enhanced cell migration and provided a sustained release of EGF, which significantly promoted neural tissue regeneration. Together, these studies underscore the potential of hydrogel-based systems for controlled drug delivery to support neural repair and functional recovery after brain injury.
Yi Liu 2020 [86] (Endogenous Repair)a1)Preparation of CH hydrogel with SIPN and self-healing properties.b1)Assessment of functional recovery after implantation of CS or CH hydrogel in a TBI zebrafish model.c1)CH hydrogel alleviates brain atrophy and neurological deficits in an ICH rat model, demonstrating endogenous repair potential.
Jiaxin Zhang 2022 [211] (Drug Agent Delivery)a2)Development of in situ injectable SF hydrogel for ICH therapy.b2)Proposed mechanisms underlying the protective effects of H2S@SF hydrogel against ICH-induced cellular injury.
Parkinson's disease (PD) is the second most common neurodegenerative disorder worldwide and is characterized as a chronic, progressive condition primarily resulting from the degeneration of dopaminergic pathways in the substantia nigra and striatum [213]. Clinically, PD presents with resting tremors, impaired coordination, reduced motor function, muscle rigidity, and postural instability [214,215]. The pathological hallmark of PD is the region-specific loss of dopaminergic neurons in the substantia nigra and their projections to the striatum. This neuronal loss—especially within the compact part of the substantia nigra—disrupts normal neural circuitry by causing irregular discharges from the subthalamic nucleus, which underlies many of the disease's clinical manifestations [216,217].
Currently, PD is incurable, and existing treatments primarily aim to alleviate symptoms rather than modify the disease course. Pharmacotherapy remains the primary treatment modality, often supplemented by deep brain stimulation (DBS) and alternative management strategies. Advances in understanding the role of dopamine in PD pathophysiology have led to the development of drugs that directly or indirectly modulate dopaminergic transmission, which has significantly improved symptom management [218]. However, chronic use of these medications is frequently associated with reduced efficacy, the emergence of movement disorders, and other side effects [219]. DBS has emerged as a novel therapeutic approach, involving the implantation of electrodes to deliver electrical stimulation to specific brain regions such as the subthalamic nucleus (STN) and the globus pallidus internus (GPi) [220,221]. While DBS can effectively regulate motor function and alleviate symptoms, its impact on cognitive and psychiatric aspects of PD remains contentious, with some studies reporting progressive declines in neuropsychological performance over time [222].
Alternative management strategies for PD, including physical therapies such as tai chi, yoga, acupuncture, and dance, offer some benefits but generally yield modest improvements [223]. Therefore, there is a pressing need for continued research and the development of novel treatment approaches that more effectively address both the symptoms and the underlying neurodegenerative processes of Parkinson's disease.
Junpeng Xu et al. [224] (Fig. 7.1) developed an injectable bioactive hydrogel that leverages a novel tannic acid crosslinker to achieve endogenous repair in PD. This hydrogel was synthesized by incorporating OTA-stable gold nanoparticles containing quinone groups, which function as effective nano-crosslinkers. In vitro studies demonstrated that the hydrogel significantly promotes the proliferation and differentiation of NSCs, while its intrinsic antioxidant and anti-inflammatory properties rescued approximately 90 % of inflamed NSCs. Furthermore, when injected directly into the lesion sites of PD rats—without any additional therapeutic agents—the hydrogel exerted pronounced antioxidant and anti-inflammatory effects. This intervention alleviated irregular neural discharges in brain projection areas, enhanced motor function recovery, and reduced histological neurodegeneration. Notably, the therapeutic outcomes achieved with this bioactive hydrogel were comparable to those obtained with drug-loaded hydrogels.Fig. 7Parkinson's disease.Fig. 7
Activin B, a member of the transforming growth factor superfamily, has demonstrated significant neuroprotective effects [225]. However, its therapeutic potential is limited by a short half-life following direct injection. To address this limitation, Juan Li et al. [226] (Fig. 7.2) developed a thermosensitive injectable hydrogel utilizing multifunctional crosslinkers to enable sustained delivery of Activin B. In this system, Activin B was incorporated into poly(N-isopropylacrylamide) (PNIPAM) hydrogels, which were stereotactically injected into the striatum of a PD mouse model. This delivery system extended the release duration of Activin B to five weeks, thereby enhancing neuroprotection in the striatum. Ultimately, the PD mouse model treated with the Activin B-loaded hydrogel exhibited significant cellular protection, marked behavioral improvements, and excellent biocompatibility, highlighting the promise of hydrogel-based drug delivery systems in the treatment of neurodegenerative disorders.
Junpeng Xu 2023 [224] (Endogenous Repair)a1)Gelation mechanism of COA hydrogel.b1)Schematic of the PD rat model induced by 6-hydroxydopamine (6-OHDA). After confirming PD induction, COA hydrogel was injected into the lesion for 14 days, with therapeutic efficacy evaluated via behavioral assays, electrophysiology, and immunofluorescence, highlighting endogenous repair responses.
Juan Li 2016 [226] (Drug Agent Delivery)a2)Physical characterization of the hydrogel, including chemical structure and internal morphology of the lyophilized scaffold.b2)Injection site location in the PD rat model.c2)Hydrogel loaded with Activin B preserves striatal TH-positive dopaminergic fibers, demonstrating neuroprotective effects (microscopic images of TH-stained striatal sections).
Multiple sclerosis (MS) is the most prevalent non-traumatic disabling disease among young individuals, with its incidence rising globally and imposing a substantial socioeconomic burden [227]. MS is primarily a chronic inflammatory disorder of the central nervous system, marked by perivenous lesions that give rise to demyelinating plaques, oligodendrocyte destruction, and irreversible damage to both gray and white matter [227,228]. Although the exact etiology of MS remains unclear, it is widely believed to stem from interactions between genetic susceptibility and environmental factors [229].
Current treatment strategies for MS predominantly involve anti-inflammatory and immunosuppressive therapies, such as adrenal corticosteroids and various immunosuppressants [230,231]. Despite these interventions, therapeutic outcomes often fall short of expectations, highlighting an urgent need for novel treatment approaches. With the evolution of hydrogel technologies and their increasingly sophisticated structures and functions, these biomaterials are poised to play a critical role in developing new therapeutic strategies for MS.
Numerous clinical trials have demonstrated that stem cell infusion can markedly improve the clinical course of MS [232]. However, systemic administration of MSCs faces significant challenges, including low cell survival rates and poor penetration of the central nervous CNS [233]. To overcome these limitations, Helena Ferreira et al. [234] (Fig. 8.1) developed an injectable hydrogel composed of phospholipids and hyaluronic acid to deliver BMSCs directly into the CNS in an experimental autoimmune encephalomyelitis (EAE) rat model. Their comprehensive physicochemical characterization, alongside in vitro and in vivo assessments, revealed that the hydrogel exhibited strong affinity for the injured corpus callosum—a critical region implicated in neuronal deficits in MS. These findings suggest that such hydrogel-based cell delivery systems could significantly enhance targeted cell engraftment and therapeutic efficacy in MS treatment.Fig. 8Multiple sclerosis (MS).Fig. 8
Alzheimer's disease (AD) is the most prevalent neurodegenerative disorder, primarily characterized by progressive memory impairment and cognitive decline that eventually lead to severe, incurable dementia [235]. Patients with AD often suffer from reduced self-sufficiency, imposing a substantial burden on families, caregivers, healthcare systems, and society as a whole [236]. Pathologically, AD predominantly affects the medial temporal lobe and associative neocortical structures. The disease is marked by two key neuropathological neuroplaques and neurofibrillary tangles. Neuroplaques are largely attributed to the accumulation of amyloid-beta peptides in the brain, while neurofibrillary tangles result from hyperphosphorylation of tau proteins, leading to alterations in the neuronal cytoskeleton [237,238].
Currently, AD remains incurable, and treatment options are primarily symptomatic. In contrast to Parkinson's disease, where a variety of therapeutic strategies have been explored, available treatments for AD are limited to pharmacological interventions. The primary medications include cholinesterase inhibitors, such as donepezil, rivastigmine, and galantamine, and glutamate antagonists, such as memantine, which are typically prescribed for long-term management [239]. Cholinesterase inhibitors offer moderate cognitive benefits, whereas glutamate antagonists can significantly improve cognitive function when used as monotherapy in moderate to severe dementia cases [240,241]. A major challenge with these therapies is the need for the drugs to traverse the body and cross the BBB to reach their targets, resulting in low utilization rates and reduced efficacy. Consequently, there is a pressing need to develop more effective drug delivery methods. One promising avenue is nasal drug delivery, which allows for direct transport from the nasal cavity to the brain, thereby minimizing drug distribution to non-target sites and reducing systemic side effects. Fluid hydrogels designed for intranasal delivery have shown potential in enhancing targeting and controlling drug release. For example, Thomas Adnet et al. [242] developed a Poloxamer-based hydrogel that extends the residence time of active pharmaceutical ingredients (APIs) in the nasal cavity, offering controlled drug release profiles.
To date, few studies have investigated the application of injectable hydrogels in animal models of AD. Song Yi Lee et al. [243] (Fig. 8.2) developed a directionally double-crosslinked hydrogel using ferrous sulfate (FeSO4) for the subcutaneous delivery of donepezil (DPZ). Their findings suggest that hydrogel-assisted subcutaneous injection may reduce the frequency of administration and facilitate the efficient and safe long-term delivery of DPZ in Alzheimer's patients. Despite these advances, significant breakthroughs in the use of injectable hydrogels for AD treatment remain elusive, underscoring the need for further investigation in this area.
Helena Ferreira 2021 [234] (Exogenous Cell Delivery)a1)Cross-sectional SEM images of hydrogels composed of 1 % hyaluronic acid (HA) or 1 % HA incorporating liposomes at 7.5 mM and 15 mM, showing structural differences.b1)Biological responses of BMSCs cultured in hydrogels with varying HA and liposome cell viability, proliferation, and total protein synthesis over 1, 3, and 7 days, indicating dose-dependent enhancement (∗p < 0.01 vs. control).c1)Functional outcomes in EAE rats after ICV injection of hydrogels containing 250,000 or 750,000 BMSCs, showing improved cumulative disability and maximum clinical scores, demonstrating the therapeutic potential of exogenous cell delivery (∗p < 0.05).
Alzheimer's disease (AD)
Song Yi Lee 2020 [234] (Intranasal Administration)a2)Schematic of an in situ double-crosslinked hydrogel formed via single-syringe injection, designed for intranasal delivery.b2)Pharmacokinetic analysis in rats after subcutaneous injection of DPZ formulations (solution, HA-dp/PD MS, and HA-dp/PD MS/FeSO4), illustrating sustained drug release profiles and improved bioavailability with hydrogel systems.
SCI is among the most severe forms of trauma, typically resulting from traffic accidents, falls, or other incidents that cause contusions, compression, or transection of the spinal cord. Inflammation and tumors can also contribute to these injuries [244]. SCI leads to significant dysfunction in the sensory, motor, and autonomic nervous systems, placing substantial physical, emotional, and economic burdens on patients and their families.
SCI unfolds in distinct stages, making it critical to understand the timing, development, and underlying pathophysiology to devise effective treatments. The initial phase, known as primary injury, involves the immediate, irreversible damage inflicted by direct or indirect external forces on the spinal cord. Although brief, this phase is quickly followed by a prolonged period termed secondary injury [245]. Secondary injury can propagate damage beyond the original injury site, further impairing spinal cord function. Its pathophysiology is complex and includes inflammatory cell infiltration, axonal damage, apoptosis of neurons and glial cells, and glial scar formation [[246], [247], [248]].
Recent advancements in biomaterials and tissue engineering offer promising alternatives, particularly through the use of implantable or injectable bioengineering scaffolds. Injectable hydrogels, in particular, are gaining attention due to their favorable properties—high biocompatibility, hydrophilicity, excellent cell interaction, permeability, injectability, and biodegradability—which render them attractive candidates for SCI repair. Current therapeutic strategies for promoting neural repair and regeneration following SCI are primarily focused on three directions as following 1. Creating a supportive milieu that facilitates unrestricted nerve growth by reducing inhibitory barriers to axonal sprouting. 2. Rebuilding and repairing damaged neural tissue to restore functional connectivity. 3. Utilizing spinal cord-like tissue implants that support and guide axonal regeneration [249].
Gong Ho Han et al. [250] developed a synthesized TUDCA hydrogel (TC gel) that was soaked in a TUDCA solution for 1 h before being injected into a rat model of SCI. Their findings demonstrated that the TC gel exerted a significant inhibitory effect on neuroinflammation. In a contrasting approach, Luzhong Zhang et al. [251] employed a straightforward in situ Michael addition method to synthesize an ibuprofen-kyigsrk conjugated hydrogel—without the need for catalysts or additional reaction steps. This hydrogel was shown to markedly inhibit inflammatory responses while promoting nerve regeneration. Differing from these methods, Shangzhi Li et al. [252] (Fig. 9.1) developed a hyaluronic acid-based hydrogel, termed AHA/DTP. In a rat model of SCI, the AHA/DTP hydrogel effectively bridged the lesion site, facilitating angiogenesis, myelin regeneration, and nerve repair, which culminated in significant motor function recovery.Fig. 9Spinal cord injury.Fig. 9
Hydrogels have been extensively employed as delivery vehicles for exogenous cells in SCI therapy. These hydrogels are loaded with various cell types—most notably NSCs—which, once injected into the injury site, differentiate into neuronal and glial cells. This differentiation facilitates the secretion of cytokines that suppress inflammation and apoptosis, promote axonal regeneration, and help restore inter-neuronal communication. Numerous studies by Shangzhi Li et al. [157], Yian Luo et al. [253], Zhiping Qi et al. [254], Dun Liu et al. [255], and Hou Liu et al. [256] have demonstrated that NSCs loaded within hydrogels enhance neuronal differentiation, axonal growth, and myelin regeneration, ultimately aiding in motor function recovery. In particular, Dun Liu et al. (Fig. 9.2) employed a CeNP-Gel hydrogel in a rat contusion model. This hydrogel not only reduced oxidative stress in the injury microenvironment but also promoted the survival, integration, and neural differentiation of encapsulated NSCs, thereby mitigating glial scar formation and enhancing neurogenesis. Similarly, Hou Liu et al. used an AT-OHA hydrogel in a rat contusion model; in addition to promoting axonal regeneration, reducing cavity formation, and inhibiting glial scar proliferation, their hydrogel also established an electrophysiological platform for cell communication in a complete spinal cord transection model, further accelerating tissue regeneration and motor recovery.
Other studies have utilized different hydrogel formulations for NSC delivery. Yian Luo et al. synthesized a BOCPG hydrogel through Schiff base and non-covalent crosslinking in a rat hemisection model, while Zhiping Qi et al. employed a cfGel hydrogel in a contusion model. Additionally, Shangzhi Li et al. developed a hyaluronic acid-based hydrogel (AHA/DTP) via Schiff base crosslinking, which has been investigated in both central and peripheral nervous system contexts.
Beyond NSCs, other cell types such as MSCs, progenitor cells, and ESCs have been explored due to their robust self-renewal and neuronal differentiation capabilities. For instance, Kristyna Zaviskova et al. [257] synthesized an HA-PH-RGD/F hydrogel using enzyme crosslinking and incorporated Wharton's jelly-derived MSCs. When injected into a rat spinal cord hemisection model, the hydrogel bridged the lesion cavity, facilitated vascularization, and promoted axonal sprouting after eight weeks. Similarly, Christy Kwokdinata et al. [258] demonstrated that a hyaluronic acid-gelatin hydrogel encapsulating spinal cord progenitor cells improved cell viability and localization during in vivo implantation. Chong Wang et al. [259] utilized a GCP-hydrogel to deliver ESCs in a rat model with a quarter-section SCI, showing that the composite hydrogel significantly enhanced tissue regeneration, vascularization, and the expression of neural markers, like Tuj1, MAP2, and Syn, while also reducing immune responses. Finally, Schwann cells—essential glial cells in the peripheral nervous system—have been encapsulated within hydrogels to further support regeneration. Laura M Marquardt et al. [260] developed the SHIELD hydrogel using Michael addition to encapsulate Schwann cells. Their study demonstrated that the SHIELD hydrogel effectively reduced neuronal damage, secondary injury, and cystic cavitation, offering a novel therapeutic approach for SCI treatment.
Systemic administration of many drugs fails to deliver sufficient therapeutic concentrations to the site of SCI due to the restrictive nature of the blood-spinal cord barrier. Injectable hydrogels offer an attractive solution by creating a local microenvironment conducive to cell adhesion and growth while allowing the sustained transport of essential nutrients and therapeutic agents. This targeted delivery approach not only enhances drug retention at the injury site but also improves cellular survival and tissue regeneration. For the anti-inflammatory agents, diacerein, as an interleukin-1β inhibitor with well-documented anti-inflammatory properties, diacerein has been repurposed for SCI repair. Kaijia Zhang et al. [261] employed PD/GO hydrogels to deliver diacerein, demonstrating significant modulation of astrocyte proliferation in vitro and effective in vivo repair of SCI. Although primarily used for Alzheimer's disease, donepezil has been incorporated into sodium hyaluronate oxide (SAO) and polyaniline-grafted gelatin (NH2-Gel-PANI) hydrogels by Tiemei Liu et al. [262] The resulting hydrogels exhibit excellent sustained release properties that promote neural stem cell differentiation and foster new tissue formation.
For the ionic and small molecule therapeutics, Lithium chloride (LiCl) is widely used as a mood stabilizer for bipolar disorder. Fang Wang et al. [263] (Fig. 9.3) incorporated LiCl into PLL hydrogels for direct injection into a spinal cord hemisection model. The findings indicated that LiCl enhances electrical signal transmission, reduces oxidative stress, modulates inflammation, promotes axon growth, and reconstructs synaptic structures, thereby accelerating post-surgical repair and functional recovery in spinal cord injuries. Another strategy is zinc supplementation, typically, SCI is linked to Zn^2+^ deficiency, a major cause of glutamate excitotoxicity and local neuronal death. Zeolitic imidazolate framework-8 (ZIF-8) is commonly employed as a drug that continuously releases Zn2+ in acidic environments. Heng Zhou et al. [264] incorporated ZIF-8 and dental pulp stem cells (DPSCs) into gelatin methacryloyl (GelMA) hydrogels. Following injection into the spinal cord injury site, ZIF-8 compensated for Zn2+ loss and promoted the neurodifferentiation of DPSCs and the secretion of VEGF-A, facilitating vascular restoration and nerve regeneration. For the anti-fibrotic and scar-inhibitory agents, decorin as a naturally occurring anti-inflammatory molecule and TGFβ1/2 antagonist, decorin exhibits multiple regulatory effects on the central nervous system, including the inhibition of scar formation and suppression of fibrosis. Jacob Matthews et al. [265] utilized BC/FB hydrogels to deliver decorin, resulting in improved electrophysiological performance, motor and sensory function recovery, and overall neural repair.
Furthermore, Studies have demonstrated that exosomes (Exos) derived from MSCs can modulate paracrine mechanisms to enhance tissue repair. Jiyun Cheng et al. [266] developed a GelMA hydrogel system to deliver bone MSC-derived exosomes, which promoted Tuj-1 positive neuronal differentiation, reduced astrocytic scarring, and increased axonal elongation. On the other hand, Harnessing the natural drug-carrying capacity of small extracellular vesicles (sEVs) derived from human umbilical cord MSCs (hUC-MSCs), Heng Wang et al. [267] employed photopolymerization to fabricate a GelMA hydrogel incorporating sEVs loaded with berberine (sEVs-Ber). In a rat hemisection model, this dual-delivery system reduced local inflammation and injury severity, thereby creating favorable conditions for nerve regeneration and axonal growth, which ultimately facilitated motor function recovery.
Lastly, as we know, neurotrophic factors are widely recognized for their ability to promote neuronal survival and axon growth during embryonic development, protect damaged neurons, and enhance the proliferation and differentiation of NSCs. As a result, numerous studies have integrated growth factors into injectable hydrogels, allowing these factors to cross the blood-brain barrier directly to the damaged targets, while mitigating risks linked to long-term use. In 2022, Linquan Zhou et al. [268] developed a methacrylate-silk fibroin hydrogel through photo-crosslinking, which was loaded with basic fibroblast growth factor (bFGF). Final in vitro and in vivo experiments demonstrated that sustained release of bFGF from the hydrogel improved mitochondrial function, inhibited glial cell proliferation, and promoted axon regeneration in injured neurons, indicating a promising direction for spinal cord repair.
Shangzhi Li 2022 [157] (Endogenous Repair)a1)Schematic of the synthesis pathway for the injectable AHA/DTP hydrogel.b1)In situ injection of hydrogel into the SCI lesion site.c1)Self-repair mechanism mediated by the hydrogel.d1)Histological analysis of SCI rats treated with AHA/DTP hydrogel at 56 days post-injury; spinal cord sections stained for CD31 (red) and DAPI (blue), showing enhanced vascularization and tissue recovery.
Dun Liu 2023 [269] (Exogenous Cell Delivery)a2)Preparation of HSP-F/BCS hydrogel and its application in the SCI model.b2)Dual immunofluorescence staining for NF200 (red), MBP (green), and DAPI (blue), highlighting axonal regeneration and myelination; higher magnification indicated by white boxes.c2)LFB staining of injured spinal cords demonstrating myelin preservation and tissue repair under hydrogel treatment; magnified regions shown in black boxes.
Fang Wang 2023 [263] Drug/Biological Agent Delivery.a3)Schematic of PLL hydrogel preparation and its potential application in SCI repair.b3)HE and LFB staining combined with electrophysiological assessment, showing structural and functional improvements at the injury site under hydrogel-mediated drug delivery.
Peripheral nerves are delicate structures forming a complex, organized network that connects the CNS to distal target organs involved in motor and sensory pathways [270]. The PNS is composed of two main the parenchyma and the stroma. The parenchyma consists of nerve fibers, predominantly axons enveloped by Schwann cells, while the stroma is made up of specialized connective tissue [271]. Due to their extensive distribution, peripheral nerves are highly vulnerable to traumatic injuries from various anatomical sites, and they can also be affected by tumors and other pathological conditions.
Peripheral nerve injury (PNI) is a complex and frequently encountered condition, particularly in younger individuals. It can result in temporary or permanent loss of sensory and motor functions, accompanied by persistent neuropathic pain. PNI is associated with high incidence rates, limited treatment options, and generally poor clinical outcomes [272,273]. The condition not only leads to functional impairments and psychological distress but also restricts daily activities, professional engagement, and recreational pursuits, thereby imposing a significant societal burden [274].
For traumatic PNI, surgical treatment is typically the primary approach. Common surgical techniques include nerve release, end-to-end neurorrhaphy, nerve grafting, and side-to-side neurorrhaphy. In cases where large nerve gaps are present, autologous nerve grafting is considered the gold standard for repair [275,276]. However, the success rate of autologous nerve grafting remains relatively low [277], and the technique is associated with several drawbacks, such as extended surgical times, the potential need for multiple procedures, and donor site complications including painful neuromas, scarring, and sensory loss [278,279]. To overcome these limitations, microtube technology has emerged as a viable alternative for bridging nerve gaps [280]. Compared to traditional surgical methods, nerve conduits offer advantages such as shorter operative times, reduced suture requirements, and the avoidance of secondary interventions and donor site complications [281,282]. Nonetheless, conventional nerve conduits still face significant challenges, including rigidity, poor flexibility, and incompatibility with nerve tissue [[283], [284], [285]].
Recent advancements in tissue engineering have prompted researchers to develop innovative treatment strategies for peripheral nerve regeneration. Among these, injectable hydrogels show significant promise due to their excellent biocompatibility, tunable mechanical properties, and ability to mimic the natural extracellular matrix. These hydrogels can be administered via minimally invasive methods, offering a novel approach to bridge nerve gaps and enhance functional recovery.
Zhenwei Yi et al. [286] (Fig. 10.1) developed a dual-crosslinked hydrogel consisting of carboxymethyl chitosan (CMCS), HA, and polyaniline (PANI), which exhibits injectable, self-healing, and conductive properties. The study found that injecting the self-healing hydrogel, specifically the 3 % concentration of ALHA/CMCS/CP (ACCP), at the site of sciatic nerve compression injury reduced tissue electrical resistance, increased nerve conduction velocity, decreased the sciatic nerve functional index, enhanced the expression of neuron-specific axonal proteins, induced myelination, and prevented muscle denervation atrophy. Hongyun Xuan et al. [287] developed HASPy hydrogel, which directly targets interleukin-17 receptor A (IL-17RA) to promote the expression of genes and proteins associated with Schwann cell myelination via activation of the interleukin-17 (IL-17) signaling pathway. The hydrogel was injected directly into the site of sciatic nerve compression injury in rats to assess its in vivo nerve regeneration capability, revealing that it facilitated functional recovery and myelin regeneration. In contrast, Travis A Prest et al. [288] developed a decellularized hydrogel derived from peripheral nerve decellularization and nerve-specific components. The study demonstrated that injecting the nerve-specific hydrogel directly into the injured nerve did not disrupt the nerve regeneration process and had significant potential to accelerate recovery following nerve damage.Fig. 10Peripheral nerve injury.Fig. 10
Stem cell therapy has emerged as a prominent research focus for both peripheral and central nervous system injuries. Among the various types of stem cells, mesenchymal stem cells (MSCs) have demonstrated considerable potential in repairing damaged tissues, including peripheral nerve injuries. For instance, Majid Salehi et al. [289] utilized a decellularized matrix hydrogel (DAM-gel) to deliver olfactory ensheathing mesenchymal stem cells (OE-MSCs) into a rat model of sciatic nerve compression injury. The combination of OE-MSCs with an alginate/chitosan hydrogel markedly improved recovery outcomes in the injured animals.
Research indicates that endometrial stem cells have a high potential for differentiation into neuron-like cells. Due to their strong capacity for differentiation into neural cells, Zahra Taherian Mobarakeh et al. [290] (Fig. 10.2) employed these stem cells as regenerative agents for sciatic nerve tissue repair. The developed fibrin gel, which incorporated insulin-loaded chitosan nanoparticles (Ins-CPs) and endometrial stem cells, significantly improved motor function and sensory recovery when loaded with insulin.
Numerous drugs, biological agents, ions, and molecules exhibit neuroprotective or neurorepair properties. However, because these agents cannot directly access damaged sites, employing hydrogels for direct injection has emerged as a promising strategy.
Firstly, in the context of drugs, Curcumin (Cur) is a natural polyphenolic compound extracted from turmeric rhizomes, traditionally used in herbal medicine for treating inflammation. Yunfan Kong et al. [291] utilized boronate-based hydrogels to deliver curcumin into a sciatic nerve injury model. The results indicated that this approach reduced local oxidative stress and inflammation, protected nerve structures, prevented muscle atrophy, and mitigated neuroinflammation and neuronal sensitization within the nervous system.
Secondly, these substances, known as biological agents, are typically employed to facilitate biological processes (such as angiogenesis or cellular repair) rather than being utilized directly as conventional pharmacological drugs.
Vascular endothelial growth factor (VEGF) is widely recognized for promoting angiogenesis by binding to its receptors and has well-documented neuroprotective and neurotrophic effects, including enhanced neuronal survival and axonal development [[292], [293], [294], [295]]. In one approach, Wanlin Xu et al. (Fig. 10.3) incorporated VEGF into hydrogels in both the 2023 [296] and 2021 [150] studies. The former study employed a synthetic VEGF mimetic peptide (QK), capable of binding to and activating VEGF receptors. The researchers utilized nanoliposomes (QK-NLs@Gel) for direct injection of the QK peptide into a rat model of compressed sciatic nerve, demonstrating that this hydrogel sustained the release of angiogenic QK for PNI repair. During recovery, the QK-NLs@Gel hydrogel enhanced vascular reconstruction and M2 polarization of macrophages, thereby facilitating and accelerating nerve regeneration. In contrast, the latter study integrated VEGF into a phot cross-linked GelMA hydrogel and applied it to a sciatic nerve compression injury model. The study found that the controlled release of VEGF from the composite hydrogel accelerated axonal regeneration, myelin sheath repair, and vascular reconstruction in the crushed sciatic nerve. This suggests that injectable hydrogels delivering VEGF offer new options for the clinical treatment of peripheral nerve compression injuries.
In addition to VEGF, Exosomes, which transport genetic material and proteins, play a critical role in modulating recipient cell functions, promoting neuronal proliferation and migration at injury sites, and enhancing overall tissue repair. Consequently, exosomes derived from mesenchymal stem cells (MSCs) have emerged as a novel therapeutic strategy. Xiaodie Ma et al. [297] explored the potential of injectable hydrogels as delivery vehicles for MSC-derived exosomes to treat diabetic peripheral neuropathy (DPN). Given that DPN typically requires prolonged treatment and that the effects of exosomes manifest slowly—necessitating repeated administrations—the authors developed a graphene oxide–polymer hybrid hydrogel cross-linked via Schiff bases to incorporate stem cell-derived exosomes. In vitro experiments revealed that the synergistic combination of the GC-PHOM hybrid hydrogel and exosomes significantly enhanced biocompatibility and promoted Schwann cell growth and migration, suggesting promising applications for both DPN treatment and nerve repair. Similarly, Shuai Liu et al. [298] investigated the use of HG hydrogels to deliver exosomes derived from adipose-derived MSCs (ADSCExo) for the treatment of cavernous nerve injury associated with erectile dysfunction (CNI-ED), a common complication following radical pelvic surgery. In a CN/pelvic ganglia (MPG) injury model, treatment with HG@Exo significantly improved erectile function, a benefit attributed to enhanced Schwann cell viability and increased neurite outgrowth.
Aminopyridine (4-AP), is a broad-spectrum potassium channel blocker known to enhance neuromuscular function in patients with various demyelinating diseases. Saba Nemati Mahand et al. [299] utilized photo-crosslinked methacrylate (GelMA)/poly(2-ethyl-2-oxazoline) (PEtOx) hydrogels for the delivery of 4-AP. Cells within the GelMA/PEtOx hydrogel demonstrated excellent proliferation capabilities, indicating potential for promoting nerve regeneration.
Bacteriophages are viruses that utilize bacteria as hosts. Yingfan Chen et al. [300] aimed to utilize r-phage as a functional material for repairing extensive nerve defects. They combined fibrin gel (FG) with bacteriophages and incorporated neural stem cells, subsequently injecting this mixture into a rat model of a 10 mm sciatic nerve gap. They found that this resulted in a solid hydrogel fiber rich in neurons, functioning as an artificial nerve conduit, which stimulated neurogenesis at the graft site within 60 days, facilitating nerve regeneration.
Insulin, an important signaling molecule, is widely used in various tissues to address functional defects and regulate cellular metabolism. Vahid Zolfagharzadeh et al. [301] utilized HAPh/ColPh microspheres to deliver insulin, which were directly injected into a rat model of sciatic nerve compression. The results indicated that the biomimetic microspheres promoted the regeneration of sciatic nerve tissue, especially when loaded with insulin. These findings suggest that the generated biomimetic microspheres, loaded with bioactive molecules, provide a useful platform for regulating natural healing and inflammatory responses in damaged or diseased sciatic nerve tissue.
Brachial plexus avulsion (BPA) is a severe traumatic peripheral nerve injury that leads to upper limb paralysis. Kangzhen Chen et al. [302] employed Pluronic F-127 (PF-127) to deliver miR-615, a potential microRNA (miRNA) that negatively regulates LINGO-1. They created a BPA model and injected the miR-615-loaded hydrogel, which suppressed astrocyte activation and promoted functional recovery after BPA, facilitating neurogenesis. This study indicates that miR-615 may serve as a therapeutic target for BPA treatment.
Furthermore, ions, which are substances primarily involved in cellular signaling or biochemical reactions, are commonly utilized to stimulate physiological responses in organisms.
Zhi Yao et al. [303] incorporated bisphosphonates to deliver magnesium (Mg), whereby Mg^2+^ promoted neurite growth in a concentration-dependent manner by activating the PI3K/Akt signaling pathway and Sema5b. In a 10 mm sciatic nerve defect model, the HA-Pam-Mg hydrogel exhibited significant shear-thinning properties and injectability. Over time, this hydrogel, combined with 3D-engineered PCL conduits, facilitated peripheral nerve regeneration and functional recovery.
Finally, molecules, including those with specific bioactive properties such as berberine, although it is a natural compound, function not merely as pharmaceutical agents but rather as multifunctional entities with diverse biological roles.
Berberine is an alkaloid extracted from the rhizomes, roots, and stems of several plants. Research has demonstrated the antibiotic, antitumor, and anti-peristaltic properties of berberine. Majid Rahmati et al. [304] utilized chitosan (Chit) in combination with alginate to deliver berberine into a sciatic nerve compression model. The results indicated that the prepared Alg/Chit hydrogels containing different concentrations of berberine exhibited a structure suitable for nerve tissue engineering, with both cell compatibility and blood compatibility, demonstrating a positive effect on the healing of sciatic nerve injuries.
Zhenwei Yi 2023 [286] (Endogenous Repair)a1)Schematic of injectable conductive hydrogel preparation and application in rat PNI.b1)Morphometric analysis of gastrocnemius muscle demonstrating muscle recovery on injured versus healthy sides.c1)TB staining of distal nerve cross-sections at 4 and 8 weeks, highlighting axonal regeneration and myelination.
Zahra Taherian Mobarakeh 2023 [305] (Exogenous Cell Delivery)a2)Surgical schematic for injection of fibrin hydrogel to repair rat sciatic nerve injury.b2)Hydrogel containing Ins-CP and hEnSC bridges a 4 mm sciatic nerve defect, supporting long-term axonal regeneration and functional recovery.
Wanlin Xu 2023 [306] (Drug Agent Delivery)a3)Schematic of injectable nanoliposome-encapsulated VEGF-mimetic peptide (QK-NLs@Gel) hydrogel for PNI repair.b3)TEM imaging of regenerated nerve fibers at day 28 post-surgery, showing enhanced structural recovery.c3)Muscle functional recovery at day 28 post-surgery, illustrated by gastrocnemius and soleus muscle morphology.
This review details the standards for biomaterials that facilitate nerve regeneration, the types of injectable hydrogels created by various preparation methods, and recent applications of biomaterials in nerve tissue engineering. In conclusion, we describe the applications of injectable hydrogels in the nervous system as "flexibility."
On one hand, the injectable characteristics enable administration in various regions of the brain, spinal cord, and peripheral nerves, allowing for precise filling of irregularly shaped lesions through in situ molding. Additionally, beyond the inherent therapeutic properties of the hydrogel, it can effectively transport drugs, therapeutic molecules, and cells, thereby enhancing their efficacy.
Neurobiologically, the CNS and PNS exhibit distinct barrier mechanisms and repair processes following injury. The CNS is protected by the BBB, BCSFB, and BSCB, complicating the delivery of traditional therapies to damaged sites, while the PNS has its own protective mechanisms. Despite these differences, both systems face similar the need for minimally invasive methods to effectively reach and treat lesions. Injectable hydrogels provide an optimal solution, as they can be administered via syringe-sized devices and, being liquid, conform to irregularly shaped injury sites.
Pathophysiologically, the injury response in the PNS is relatively straightforward, primarily involving Schwann cell activity and Wallerian degeneration. In contrast, CNS tissues have a limited regenerative capacity, largely due to post-injury responses such as neuroinflammation, glial scarring, and neurodegeneration. Nonetheless, injectable hydrogels can meet the therapeutic demands of both systems by delivering drugs, bioactive molecules, and cells, enabling tailored interventions.
Both the CNS and PNS have unique material selections for injectable hydrogels, with both natural and synthetic polymers being widely applicable in both systems. For example, GelMA has been utilized to prepare injectable hydrogels for applications in the brain, spinal cord, and peripheral nerve [150,151,266]. In addition, materials such as HA [86,157,307] and collagen [187,269,308] have also been extensively employed in injectable hydrogel applications across the brain, spinal cord, and peripheral nerves. Spearman et al. (2020) quantified the elastic modulus of multiple rat soft tissues, including the tibialis anterior (TA) muscle, sciatic nerve, kidney, spleen, heart, liver, brain, spinal cord, and lung. The TA muscle exhibited the highest modulus (4.28 ± 0.51 kPa), whereas the lung showed the lowest (0.31 ± 0.04 kPa). Importantly, their findings revealed significant inter-tissue differences in mechanical modulus among the sciatic nerve, brain, and spinal cord [309]. This raises the Can a single injectable hydrogel material be suitable for all three tissue types? The answer is negative. Numerous studies have shown that the mechanical modulus of hydrogels can be modified by varying material concentrations to align with the target tissue. For instance, Yan Hu et al. [180] employed dynamic boronate ester cross-linked hydrogels created from phenylboronic acid-grafted hyaluronic acid (HA-PBA) and dopamine-grafted gelatin (Gel-Dopa). By varying the concentration of Gel-Dopa, they obtained injectable hydrogels with different mechanical moduli (3 % Gel-Dopa: 30.3 ± 10.5 Pa; 4 % Gel-Dopa: 157 ± 74.5 Pa; 5 % Gel-Dopa: 589 ± 98.2 Pa). This indicates that within the same injectable hydrogel material, desired characteristics can be attained through personalized adjustments, demonstrating the flexibility of these hydrogels.
This article outlines the prospects for developing injectable hydrogels. Numerous therapeutic approaches in animal models have shown the ability to prevent further neural tissue damage and effectively promote nerve regeneration; however, similar results have not been observed in clinical applications. This discrepancy arises because nerve injury is multifaceted and involves continuous pathological processes. Thus, a comprehensive treatment strategy is essential for managing neural injuries. Future therapeutic strategies should not focus solely on one or two pathophysiological aspects; instead, a multifaceted treatment incorporating various therapies will be necessary for improved clinical application.
However, it is important to acknowledge that not all applications of injectable hydrogels in neural tissue repair have been successful. Several studies have reported limited therapeutic outcomes or even adverse effects. For example, PEG-based hydrogels with rapid degradation rates failed to provide sufficient mechanical support for axonal growth, while alginate or chitosan hydrogels sometimes induced fibrotic encapsulation and inflammatory responses, ultimately impeding neural regeneration [105,310]. In addition, inappropriate stiffness or mismatched viscoelasticity between the hydrogel and host tissue can negatively influence cellular behaviors such as migration, differentiation, and synaptic reconnection [311]. These failure cases emphasize that hydrogel design must consider not only biocompatibility and degradability but also the dynamic interaction between material properties and the neural microenvironment. Learning from these unsuccessful attempts can guide future optimization of injectable hydrogel systems for more consistent and reliable outcomes in neural repair.
In addition to the injectable hydrogel discussed above, several emerging technologies hold great promise for advancing neural repair. Stimuli-responsive “smart” hydrogels, which respond to environmental cues such as pH, temperature, or magnetic fields, can provide on-demand drug release and adaptive mechanical properties, making them particularly suitable for the dynamic microenvironment of neural injury sites [44,312]. Injectable bicontinuous hydrogels with interpenetrating dual-phase networks better recapitulate the heterogeneous architecture of native neural tissue, potentially enhancing both mechanical stability and intercellular communication [313,314]. Composite hydrogels incorporating iPSC-derived neural stem cells offer a patient-specific approach to restore lost neural tissue and function [315,316]. However, challenges such as long-term biocompatibility, precise spatiotemporal control of bioactive release, and large-scale clinical translation remain to be addressed. The continued integration of these innovative strategies may guide the development of the next generation of injectable hydrogels for CNS and PNS repair.
Current experimental animal models primarily consist of mammals, such as rats and rabbits, with limited focus on regenerative models like salamanders, clawed frogs, and planarians. Whether concerning the CNS or PNS, the essence of repair is the regeneration of neural cells, which regenerative models exemplify effectively. Slater et al. [317] found that, in regenerative models, apoptosis is essential for inducing cell proliferation and regeneration, with oxidative reactions and cell death also playing critical roles. This supports the idea that, in regenerative organisms, the injury itself can initiate regenerative responses. Consequently, in-depth research on regenerative models may reveal unexpected therapeutic strategies for adaptive regeneration after nerve injury. In the future, incorporating specific characteristic factors or cells from certain regenerative models into injectable hydrogels for direct application to damaged neural cells in mammals may improve therapeutic outcomes. In addition to in vivo studies, injectable hydrogels have shown promise in in vitro models of both the CNS and PNS. These systems offer a controllable microenvironment to study neural cell behaviors—such as proliferation, differentiation, and neurite outgrowth—under defined biochemical and mechanical conditions. For instance, hydrogels composed of GelMA, collagen, or hyaluronic acid can effectively mimic the extracellular matrix, facilitate neural network formation, and serve as delivery platforms for growth factors or stem cells. Integrating discoveries from in vivo regenerative models with insights gained from in vitro hydrogel systems may provide valuable guidance for the rational design of next-generation injectable scaffolds for neural repair.
Despite the favorable effects of injectable hydrogels demonstrated in numerous animal models—such as promoting axonal regeneration, immunomodulation, and angiogenesis—their clinical translation for neural repair remains in its infancy. According to currently available clinical trial records and published data, no completed and published randomized controlled trials have yet provided definitive evidence that injectable hydrogels can promote central or peripheral nerve regeneration and significantly restore neurological function in humans. In contrast, injectable hydrogels have already been applied in practical clinical settings for facial correction, tissue repair and regeneration, and drug delivery formulations [318]. This discrepancy primarily stems from the complexity of the neural microenvironment, which demands precise regulation of hydrogel degradation kinetics, mechanical properties, and bioactive factor release to ensure therapeutic consistency. Several reviews have summarized the current translational barriers—including quality control, long-term biocompatibility, GMP-compliant manufacturing, and regulatory pathways [17,319]. Future research should focus on improving the reproducibility of injectable hydrogel formulations, integrating multimodal therapeutic strategies (e.g., combining electrical stimulation with biochemical or cellular signaling), and establishing standardized evaluation systems to accelerate clinical approval. With the continued advancement of interdisciplinary collaboration, injectable hydrogels hold great promise for transforming from laboratory research into viable clinical therapies for neural repair.
In summary, injectable hydrogels offer unique advantages for both the CNS, which includes the brain and spinal cord, and the PNS. Different diseases impacting each system involve distinct pathologies, and even the same disease can exhibit different pathological progressions at various stages. Phan et al. describe recovery from brain strokes as a long-term dynamic process, both spatially and temporally [320]. The diverse properties of injectable hydrogels enable customized applications at any stage of disease progression. Nonetheless, our understanding of the pathological processes underlying many neurological disorders remains insufficient. Additionally, there is currently no established standard for designing injectable hydrogels for neural tissue engineering. These factors may represent opportunities for further exploration in future neuroengineering research.
Kun Wu: Writing – original draft, Supervision. Zhihe Yun: Formal analysis, Data curation. Wu Xue: Data curation, Conceptualization. Tao Yu: Data curation, Conceptualization. Anyuan Dai: Data curation, Conceptualization. Inbo Han: Validation, Supervision. Vit Kotheeranurak: Investigation, Formal analysis, Data curation. Worawat Limthongkul: Project administration, Methodology. Yanting Liu: Formal analysis, Data curation, Conceptualization. Qinyi Liu: Funding acquisition, Formal analysis, Data curation.
The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.