Many investigations have been conducted to explore new therapeutic approaches for cartilage tissue regeneration owing to the limited repair ability of cartilage. Delivery of therapeutic biomolecules is a promising strategy for enhancing tissue regeneration. However, the direct delivery of therapeutic biomolecules in vivo generally leads to a rapid loss of product or the possibility of spreading to nontarget sites, which may weaken their therapeutic effectiveness and impair the translation of these strategies into clinical practice. Recently, new tissue engineering strategies using controlled delivery systems have been considered to overcome these weaknesses by regulating the spatiotemporal distribution of the therapeutic biomolecules. Hydrogels are promising carriers for the development of biomolecule delivery systems because they are highly tunable and can be fabricated to encapsulate bioactive biomolecules and control their release. Moreover, their favorable biocompatibility and ability to integrate into host tissue may facilitate their use in tissue regeneration in a wide variety of situations. This review summarizes recent strategies for the design of hydrogels as biomolecule delivery systems for cartilage regeneration.

1. Introduction

Cartilage is the load-bearing tissue of the body. Trauma and aging often damage or disrupt the structure and function of cartilage, and even minor defects can result in mechanical joint instability or progressive damage. The repair and regeneration of articular cartilage tissue remains an open challenge owing to its avascular and alymphatic nature [14]. Current therapeutic approaches for repairing damaged cartilage include autologous chondrocyte implantation, arthroscopic microfracture, and open procedures such as osteotomy and arthroplasty. However, the repaired cartilage does not have the structure of native cartilage, and these procedures generally result in the formation of fibrocartilage and continued pain and discomfort [5]. To address this, cartilage tissue engineering (CTE) approaches that aim to develop constructs that regenerate healthy and fully functional cartilage are being actively investigated [6]. Ideally, these tissue-engineered constructs should generate neocartilage that is biochemically and biomechanically identical to native cartilage.

The ideal CTE scaffold should possess the following characteristics: high biodegradability, adhesion to host tissue, and good nutrient permeability. Polymeric hydrogels are promising as scaffolding because of their highly hydrated nature, tunable mechanical properties, good biocompatibility, controllable degradation rate [7], and ability to encapsulate cells and biomolecules [8]. Additionally, the spatiotemporal organization of therapeutic biological signals is important for tissue and organ regeneration; to address this, numerous studies have developed biomaterial-based delivery systems to regulate cell behavior by topically and continuously delivering therapeutic biomolecules to a target location [9, 10]. Polymeric hydrogels are also promising as drug delivery platforms for CTE because they can be designed to provide a physiological microenvironment for the encapsulation of therapeutic biomolecules. In addition, stimuli-responsive hydrogels, which contain stimuli-responsive units that exhibit a phase transition ability in response to stimuli [11], have been explored for controlled drug release [1214]. Stimuli-responsive hydrogels show stimuli-induced structural and volume transitions, thus providing several multidimensional applications in regenerative medicine [15]. Therefore, this review addresses recent advances in hydrogel scaffolds as therapeutic biomolecule delivery platforms for cartilage regeneration.

2. Design Specifications of Hydrogel

The use of hydrogels as drug delivery platforms for CTE should satisfy several specific requirements: 1. the degradation of hydrogels should occur in a controlled manner without cytotoxic by-products; 2. the inner pores of hydrogels should be of an appropriate size and density to ensure a high diffusion and binding of labile biomolecules throughout the scaffold; 3. the release of biomolecules should be controllable and able to be delivered to targeted cells; 4. the hydrogels should support the viability, proliferation, and chondrocytic phenotype of the encapsulated cells; 5. the hydrogels should be able to integrate with the surrounding host cartilage tissue; and 6. the hydrogels should be able to be delivered in a minimally invasive manner (e.g., injection).

Hydrogels can be synthesized using natural, synthetic, or hybrid polymer backbones. Natural hydrogels have favorable biodegradability and biocompatibility, with properties similar to those of natural extracellular matrix (ECM). The most widely used natural hydrogels for CTE include alginate [1619], hyaluronic acid (HA) [2023], and fibrin [2426]. The disadvantages of natural hydrogels include inconsistent hydration and poor mechanical properties. Additionally, the widespread application of these natural polymers is limited by their low production and complex purification from organisms. Furthermore, their poor flexibility often restricts the tailoring extent of hydrogel properties.

The properties of synthetic hydrogels are more tailorable [27]. The most widely used synthetic hydrogels for CTE include poly(lactide-co-glycolide) (PLGA) [28, 29], polyethylene glycol (PEG) [30], and poly((meth)acrylic acid) p((M)AA) [31]. Characteristic advantages of these synthetic polymers are their broad modification possibilities such as stiffness, porosity, mesh size, and degradation profiles. These properties are important for drug delivery, as the binding and controlled release of biomolecules depend on the structural properties of the chosen hydrogel [3234]. Moreover, synthetic hydrogels can be designed to be responsive to certain stimuli (e.g., ionic concentration, temperature, and pH), which is beneficial for CTE. However, synthetic hydrogels lack the inherent biocompatibility and cell bioactivity properties compared with natural polymers.

Hybrid hydrogels are complexes composed of two or more natural or synthetic polymers, which often compensate for the shortcomings of both [35]. In general, the components of a hybrid hydrogel are interconnected through physical or chemical means such as electrostatic/ionic interactions [36], crosslinking [37], or irradiation [38, 39]. The combination of natural and synthetic polymers can create hydrogel scaffolds with good physiochemical properties and biocompatibility, making them more suitable for use in drug delivery [40]. For example, some previous studies have reported that synthetic polymers have problems with initial burst release of drugs [41], but the incorporation of natural polymers improves control over the release of biomolecules [42].

Pore size and interconnectivity are critical parameters that determine scaffold performance in CTE. Cell adhesion, viability, migration, and morphology are strongly correlated with the scaffold pore size [43, 44]. Porosity and pore structure also have a strong influence on nutrient and waste transference [43, 45]. Hydrogels are attractive candidates for designing cell-loading scaffolds because of their highly tunable properties. Although extensive research studies have been carried out on the influence of hydrogel scaffold porosity and pore size on their effectiveness as CTE tools [4651], their optimal porosity and pore size for chondrocyte adhesion, proliferation, and differentiation have not been standardized because of the wide variety of hydrogel materials, synthesis techniques, seed cell sources, and application conditions. It has been reported that, when compared to materials with larger pores, hydrogel materials with smaller pores are preferable for chondrocyte proliferation and expression of cartilaginous markers [49]. For instance, Zhang et al. prepared collagen scaffolds with different ranges of pore sizes (150–250, 250–355, 355–425, and 425–500 μm) to investigate the effect of pore size on the biological behavior of seeded chondrocytes. Their results indicated that the scaffold pore size did not significantly affect the proliferation of chondrocytes; however, cartilage regeneration was obviously affected, with a diameter of 150–250 μm exhibiting the best improvement in the expression and production of type II collagen (Col II) and aggrecan [47]. Similarly, Nava et al. found that a poly(L-lactide-co-trimethylene carbonate) scaffold with 175 μm pore size and 87% porosity was suitable for the proliferation of chondrocytes, whereas scaffolds with a smaller pore size (75 μm) and porosity (75%) were more beneficial to cell metabolic activity and the synthesis of cartilaginous matrix proteins [48]. However, some studies have shown that scaffolds with larger pore sizes improve chondrogenesis. Lien et al. fabricated a series of gelatin scaffolds with pore sizes ranging from 50 to 500 μm by varying the crosslinking temperature (10–25°C). In vitro experiments showed that chondrocytes in the scaffolds with larger pores (250–500 µm) exhibited an increased proliferation and ECM secretion, whereas the cells in scaffolds with smaller pores often dedifferentiated [52]. Therefore, the pore size of cell-loaded hydrogels should be adjusted according to the conditions they will be operating under for better cartilage tissue regeneration results.

3. General Synthesis Schemes of Hydrogels

Hydrogels are mainly fabricated by the formation of crosslinked polymeric networks via physical or chemical processes. Physical crosslinking is normally achieved via methods such as ionic/electrostatic interactions [53, 54], hydrophobic interactions [55], crystallization [56], polymer chain complexation [57], and hydrogen bonding [58]. The main advantage of physical crosslinking is biomedical safety owing to the absence of chemical agents during gelation [59]. Because physically crosslinked hydrogels do not form covalent bonds, their synthesis is reversible, but they are generally unstable under physiological conditions. Chemical crosslinking strategies include enzyme-induced crosslinking [60, 61], photopolymerization [62], and “click” chemistry reactions [63]. Unlike physically crosslinked hydrogels, chemically crosslinked hydrogels form covalent bonds; the synthesis of chemically crosslinked hydrogels is therefore irreversible. They are also more stable under physiological conditions and possess better mechanical strength. However, the use of organic catalysts and solvents in chemical crosslinking may introduce environmental and biocompatibility problems [64]. Therefore, to minimize the limitations of physical or chemical crosslinking while combining their advantages, an increasing effort has focused on the fabrication of hydrogels using hybrid crosslinking techniques [65]. For example, Zhong et al. developed a single-network poly(acrylic acid) (PAA) hydrogel using a combination of ionic and covalent crosslinking [66], resulting in superior mechanical properties and water absorbency.

4. Signaling Biomolecules in Cartilage Regeneration

Cartilage regeneration is a developmental process controlled and coordinated by many signaling factors. These signaling molecules initiate or suppress cellular signaling pathways, which subsequently determine cell fates [6769]. Therefore, the selection of appropriate factors is critical for the design of a controlled delivery system. Currently, the most investigated signaling factors in CTE are transforming growth factor-beta (TGF-β), bone morphogenetic protein (BMP), insulin-like growth factor-1 (IGF-1), and fibroblast growth factors (FGFs) [70, 71] (Figure 1).

The TGF-β superfamily, which has over 30 members, is a group of versatile cytokines involved in the regulation of cartilage homeostasis and regeneration [72, 73]. Among all TGF-β family members, TGF-β3 isoforms are considered to have the highest chondrogenic potential [74]. Various TGF-β isoforms are activated during cartilage regeneration; these not only promote proteoglycan and Col II synthesis but also prevent the degradation of cartilage ECM [75, 76]. As a member of the TGF-β family, BMP is responsible for the regulation of cartilage growth and resorption. BMP signaling is necessary for the initiation of prechondrogenic condensation and chondrocyte maturation [77]. In particular, BMP-2 affects cellular proliferation and induces the rapid maturation of chondrocytes, BMP-4 participates in the synthesis of aggrecan and Col II, and BMP-7 protects chondrocytes by improving their anabolism and ability to repair damaged cartilage [78]. Mechanism analysis has demonstrated that TGF-β signals bind to the receptor TGF-β-RII, which recruits and phosphorylates TGF-β-RI, leading to the phosphorylation and activation of intracellular SMA- and MAD-related (SMAD) superfamily signals. Activated SMAD proteins form a heterocomplex with SMAD4 and other co-activators, then translocate to the nucleus to regulate chondrogenic gene expression [70].

IGF-1 is a hormone similar in molecular structure to that of insulin and is recognized as a key factor in the regulation of cell proliferation and inhibition of cell death. In cartilage tissue, IGF-1 signaling occurs in the synovial fluid of the joint and regulates the expression of various chondrogenic markers [79, 80]. IGF-1 may also modulate mesenchymal stem cell (MSC) chondrogenesis by inducing cellular proliferation and the expression of chondrocyte markers via the extracellular signal-related kinase 1/2 mitogen-activated protein kinase (Erk1/2 MAPK) pathway [70, 81, 82].

FGFs are a family of cell signaling proteins that participate in various physiological processes and play a crucial role in tissue development. Any dysfunction in FGFs may result in developmental defects [83]. FGFs are generally found in the ECM of cartilage. They are activated upon tissue loading and reduce aggrecanase activity [84]. It has been reported that the knockout of FGF-2 accelerates osteoarthritis, whereas the subcutaneous administration of FGF-2 suppresses osteoarthritis in a mouse model [85]. FGF-2 treatment of bone marrow-derived MSCs enhanced chondrogenic differentiation in subsequent 3-dimensional (3D) cultures [86]. Intra-articular administration of FGF-18 can alleviate cartilage degeneration and increase cartilage thickness in a rat osteoarthritis (OA) model, but it also increases synovial thickness and chondrophyte formation [87]. It has also been demonstrated that FGFs markedly enhance chondrogenic marker SOX9 expression in both chondrocytes and prechondrocytic cells via the activation of the mitogen-activated protein kinase (MAPK) pathway [88]. However, the literature regarding the role of FGFs in cartilage regeneration is limited, and the precise mechanisms remain to be defined.

5. Strategies for Controlling Delivery of Therapeutic Biomolecules from Hydrogels

Strategies to precisely deliver biomolecules locally and continuously are gaining popularity in the field of CTE [89, 90]. In the past few decades, researchers have applied a variety of techniques and scaffold materials to develop controlled therapeutic biomolecule delivery systems. However, tissue regeneration is a complex biological process that requires a precise regulation of the spatial and temporal gradients of various therapeutic biomolecules [91]. Because of this, their delivery should be spatiotemporally regulated for maximum efficacy [92]. Therefore, it is necessary to select appropriate polymer combinations and drug-loading methods and to adjust parameters such as the proportion, particle size, shape, and surface morphology of hydrogels for a specific therapeutic site [40]. The following sections discuss the recent progress in hydrogel-based therapeutic biomolecule delivery strategies for cartilage regeneration and highlight the recent developments that allow precise spatiotemporal control over biomolecule delivery in hydrogels.

6. Growth Factor Delivery Platform using Hydrogel Scaffold for Cartilage Regeneration

6.1. Direct Physical Entrapment and Chemical Conjugation of Growth Factors

The spatiotemporal organization of the growth factors in the polymeric matrix is critical for tissue regeneration, and bioactivity is affected by several parameters, including the pH, temperature, and porosity of the matrix [40]. With the deepening understanding of the signaling factors that regulate cell fate, the construction of hydrogel scaffolds used for cartilage regeneration has become increasingly sophisticated. Among all the spatiotemporal control strategies, most technical methods include physical entrapment and chemical conjugation. These approaches not only achieve dosage control but also localize growth factors and preserve protein bioactivity. Physical entrapment usually involves premixing biomolecules with gelators and then directly stabilized in hydrogel scaffolds during the gelation process [93]. The entrapped biomolecules are usually released from the hydrogel via diffusion through the gel structure and liberation as the hydrogel dissolves [94]. Because there is no direct linkage between the biomolecule and hydrogel, their release rate is usually related to temperature, PH, and the components and construction of the hydrogel. To effectively control the release of TGF-β3, Deng et al. incorporated methacrylated HA into a poly-d,l-lactic acid/polyethylene glycol/poly-d,l-lactic acid (PDLLA-PEG) hydrogel to form PDLLA-HA [42] (Figure 2(a)) and demonstrated that PDLLA-HA hydrogel constructs exhibited a continuous release of TGF-β3 for up to three weeks compared to the constructs without HA. In addition, these TGF-β3-preloaded constructs displayed dose-dependent enhancement of chondrogenic gene expression and glycosaminoglycan (GAG) production of human bone marrow mesenchymal stem cells (hBMSCs) both in vitro and in vivo. Similarly, Shen et al. demonstrated that photopolymerizable PDLLA hydrogel supported a long-term sustained release of TGF-β3 with the incorporation of graphene oxide (GO) nanosheets [97]. Alginate-based hydrogels represent promising microenvironments for cell culture and tissue engineering because of their injectability, biocompatibility, and biodegradability [98]. Moshaverinia et al. took advantage of the properties of alginates by designing a novel 3D drug delivery system loaded with a TGF-β1 ligand using an arginine-glycine-aspartic acid (RGD)-coupled alginate hydrogel for the microencapsulation of dental MSCs [99]. They found that TGF-β1 was consistently released from the alginate scaffold, and chondrogenic gene expression markers (Col II and SOX9) were observed by matrix staining after 4 weeks. This study indicates that an RGD-coupled alginate system can be used to encapsulate stem cells for cartilage regeneration. Gentile et al. developed a TGF-β1-loaded multilayer sodium alginate/gelatin hydrogel that was assembled by electrostatic attraction and hydrogen bonding [100]. Their results showed that this alginate-based scaffold consistently released TGF-β1 for up to 25 days and exhibited good cytocompatibility and chondrogenic promotion for bovine articular chondrocytes (BAC) it was adhered too.

Despite the simplicity of physical entrapment, the random diffusion of biomolecules is not conducive to the effective control of growth factor release. Compared with physical entrapment techniques, the chemical conjugation of growth factors to hydrogel scaffolds is a method that enhances the stability and persistence of growth factors [101]. As growth factors are either insoluble in or damaged by organic solvents, the immobilization of growth factors is generally performed via aqueous chemistry. Through chemical conjugation, growth factors can not only maintain their activity but also significantly extend their release over time, thereby resulting in better chondrogenesis. McCall et al. investigated the chondrogenic potency of TGF-β1-loaded PEG hydrogels on encapsulated human mesenchymal stem cells (hMSCs) [102]. In their study, TGF-β1 was first thiolated by a reaction with 2-iminothiolane and then loaded onto PEG hydrogels through a thiol-acrylate polymerization mechanism. When hMSCs were encapsulated in TGF-β1-loaded hydrogels, it was found that the hydrogel promoted chondrogenic differentiation of hMSCs at levels equal to or greater than those of the positive controls, where TGF-β1 was administered via the culture medium. A similar study reported a thiol-functionalized hyaluronic acid (HA-SH)-based hydrogel that covalently incorporated TGF-β1 was able to maintain the viability of the encapsulated MSCs and to promote the expression of chondrocyte ECM and chondrocytic genes; this improvement was positively correlated with TGF-β1 concentration [103]. When growth factors are chemically conjugated, they do not passively diffuse, and thus, their release occurs as the hydrogel degrades over its lifetime [104]. Therefore, the release rate of growth factors can be controlled more precisely by controlling the degradation rate of hydrogel scaffolds [105]. This can be achieved by causing the hydrogel to degrade in response to specific stimuli. Griffin et al. reported photodegradable polymerizable ortho-nitrobenzyl (o-NB) macromolecules with various functional groups at the benzylic position [106]. The authors conjugated TGF-β1 onto a PEG hydrogel and observed that the release of TGF-β1 could be controlled by light exposure, representing a promising method of temporally controlling drug release. Furthermore, photo-released TGF-β1 induced chondrogenic differentiation in human MSCs, similar to native TGF-β1. Choi et al. developed a strategy for sustained and slow TGF-β1 release using a blue light-reactive chitosan (MeGC)/collagen II hydrogel system, synthesized via covalent conjugation with an amine-to-sulfhydryl crosslinker, succinimidyl-4-(N-maleimidomethyl)cyclohexane-1-carboxylate (SMCC). SMCC formed an amide bond with the primary amino groups on the chitosan backbone at the one end and a thioether bond with the cysteine residue on TGF-β1 at the other end. The controlled-release hydrogel system was shown to promote the viability and chondrogenic differentiation of adipose-derived stem cells (ADSCs) in vitro and cartilage regeneration of chondral defects in vivo [95] (Figure 2(b)).

6.2. Incorporation of Preformed Growth Factor-Loaded Delivery Systems

The dispersion of bioactive molecules in the matrix via immobilization on the hydrogel scaffold by covalent bonding or electrostatic interactions is a widely used method for creating a controlled delivery hydrogel. However, although large molecules can be physically entrapped in the mesh of a hydrogel, the attachment of small molecules is unpredictable, and thus, a labile linkage between these small molecules and the hydrogel network is desirable. An increasing number of studies have been conducted on the design of scaffolds incorporating preformed molecular-loaded delivery systems, such as polymer particles [107], rather than a direct combination of hydrogel and therapeutic biomolecules. Bioactive molecules are first loaded into a hydrogel precursor solution, which is subsequently fabricated into polymer particles through techniques such as batch emulsion [108], microfluidic emulsion [109], lithography [110], electrohydrodynamic spraying [111], and mechanical fragmentation [112, 113]. These polymer particles are then incorporated into a secondary hydrogel. This incorporation strategy can not only achieve better immobilization of small biomolecules but also temporarily separate the loaded biomolecules and the seeded cells that are also encapsulated in the secondary hydrogel, allowing a more precise stimulation of the preseeded cells. Ahearne et al. developed TGF-β3-loaded gelatin microspheres, mixing them with agarose hydrogels that were preseeded with chondroprogenitor cells to fabricate a growth factor-releasing hydrogel [114]. This allowed the slow release of growth factors from microspheres within the hydrogel and enabled the establishment and maintenance of a chondrogenic phenotype in seeded cells. In another study, PLGA-based porous microspheres were introduced as novel drug carriers by Reyes et al. [115]. In their study, PLGA microspheres were fabricated by gas foaming in an acidic aqueous solution. BMP-2 and TGF-β1 were pre-encapsulated in PLGA microspheres, which were subsequently dispersed in a matrix of segmented polyurethane (SPU) to form a bilayer scaffold system. The SPU-PLGA constructs released the growth factors in a well-regulated manner and induced rapid, good-quality osteochondral repair. Recently, Lin et al. developed an injectable methoxy poly(ethylene glycol)-poly(alanine) (mPA) hydrogel incorporating rat chondrocytes and TGF-β3-laden PLGA microspheres. This composite hydrogel system exhibited a sustained release of TGF-β3 and led to a significant upregulation of chondrocyte-specific genes and production of cartilage ECM components in the encapsulated MSCs [116]. In brief, incorporating preformed growth factor delivery systems in hydrogels containing chondroprogenitor cells represents promising new treatment options for CTE.

6.3. Multiple Growth Factor-Loaded Delivery Platforms

Chondrogenesis involves the activation of and interaction between many cellular signaling pathways, and it is impossible for a single factor to completely initiate the signaling pathways for cartilage regeneration. Therefore, efforts have been made to develop drug delivery systems that are loaded with multiple growth factors to increase their therapeutic effect on cartilage regeneration. For instance, previous studies have described the dual release of IGF-1 and TGF-β1 from hydrogel scaffolds [117, 118]. The release of IGF-1 and TGF-β1 can be controlled in a sequential manner and subsequently activated the expression of chondrogenic gene in MSCs. Han et al. developed a conically graded chitosan-gelatin hydrogel/PLGA scaffold to simultaneously deliver TGF-β1 and BMP-2 for cartilage-bone interface repair [119]. This scaffold exhibited spatially and temporally controlled release of TGF-β1 and BMP-2. The TGF-β1-loaded chitosan-gelatin hydrogel and BMP-2-loaded PLGA scaffold promoted chondrogenesis and osteogenesis, respectively. Although significant progress has been made in osteochondral tissue engineering, successfully reconstructing tissues with compositions, structures, and functional properties that are similar to those of native tissues remains a challenge. Inspired by the gradient of ECM composition and fibrous collagen structure in native osteochondral tissue, Qian et al. produced poly(ε-caprolactone) and polyethylene glycol (PCEC) 3D-stratified scaffolds with superficial cartilage, deep cartilage, and subchondral bone layers, which were filled with methacrylamide hydrogels (GelMA) containing MSCs and the regional-specific growth factors BMP-7, TGF-1, and BMP-2 [96] (Figure 2(c)). The complex-induced chondrogenic and osteoblastic differentiation in MSCs promoted the expression of cellular phenotypes and matrix accumulation profiles similar to native tissues in vitro. In vivo experiments further confirmed that the fiber-reinforced and growth factor-loaded trilaminar hydrogel structure had a positive effect on osteochondral regeneration.

6.4. Platelet-Rich Plasma-Loaded Delivery Platforms

Platelet-rich plasma (PRP), a concentration of platelets, is a reservoir of multiple growth factors, such as isoforms of TGF, IGF, and platelet-derived growth factor (PDGF), and has been demonstrated to stimulate the migration, proliferation, and chondrogenic differentiation of stem cells [121]. Owing to its abundant therapeutic potential, PRP presents an alternative approach for the development of controlled-release systems [122]. Lu et al. developed a PRP-loaded genipin-gelatin/hyaluronic acid/fucoidan (GP-GLT/HA/FD) hydrogel by adding lyophilized PRP powder into a GP-GLT/HA/FD mixture before gelation, and this controlled delivery system exhibited a sustained release of PDGF instead of the burst release associated with traditional PRP gels [123]. The PRP-loaded GP-GLT/HA/FD hydrogel showed a reduced loss of chondrocytes and decreased lesion formation in vivo, indicating potential as an osteoarthritis therapy. To overcome the burst release of growth factors in traditional PRP gels, Liu et al. proposed photocrosslinkable PRP hydrogel glue (HNPRP) based on a photoinduced imine crosslinking (PIC) reaction [124] (Figure 3(a)). Exposure to light causes o-NB alcohol-modified hyaluronic acid (HA-NB) in a PRP hydrogel to form aldehyde groups, which subsequently react with the amino groups distributed on the fibrinogen of PRP, generating a hydrogel in situ. The release of PDGF, TGF-β, and FGF from this HNPRP hydrogel was approximately linear for 14 days, whereas an obvious burst release was observed in the traditional thrombin-activated PRP gel after 2 days. Functional tests confirmed that the system promoted the migration and proliferation of both BMSCs and chondrocytes and achieved better cartilage regeneration than thrombin-activated PRP gel. Platelet lysates (PLs) are derivatives of PRP, where the platelets of PRP are lysed to release their protein content [123]. Jooybar et al. designed a novel delivery system called the platelet lysate-incorporated hyaluronic acid-tyramine (HA-TA-PL) hydrogel via enzymatic crosslinking, and it showed a promotive effect on the viability, proliferation, and chondrogenesis of hMSCs [124] (Figure 3(b)). An overview of the hydrogels used as growth factor delivery platforms for cartilage regeneration is summarized in Table 1.

7. Gene Delivery Platform using Hydrogel Scaffold for Cartilage Regeneration

Growth factor-incorporated hydrogel scaffolds have been widely investigated as drug delivery systems. However, the therapeutic efficacy of the direct delivery of these growth factors is often hampered by their shortcomings, such as their transient action, short half-life, and need for high concentrations to be effective [126]. Gene therapy is considered an alternative approach that can control the expression of growth factors by delivering genetic material encoding the desired protein to the target site [127]. However, the use of genes is still limited by their low transfection efficiency, rapid degradation, and short-term transgene expression levels [128]. Incorporating genes into hydrogels may protect DNA against enzymatic degradation and increase transfection efficiency [129] (Figure 4). Therefore, gene-loaded hydrogel systems may serve as an attractive alternative in CTE.

7.1. Viral Vector-Based Gene Delivery Platforms

Successful gene therapy requires optimal gene delivery. Currently, viral and nonviral vectors are the most commonly used gene delivery vectors [130, 131]. Using viral vectors generally involves direct administration of the vectors, which allows for the controlled and minimally invasive delivery of the genes in a spatiotemporally precise manner, reducing the possible spread and loss of gene products [132]. Madry et al. incorporated a recombinant adeno-associated virus (rAAV) vector in a thermosensitive hydrogel based on poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) (PEO-PPO-PEO) poloxamers. The rAAV vector overexpressed the chondrogenic SOX9 transcription factor, and the hydrogel system controlled the release of the virus vector in a full-thickness chondral defect minipig model [129]. Comprehensive analyses indicated that the rAAV vector system significantly improved cartilage regeneration and protected the subchondral bone plate from early bone loss. Maihöfer et al. reported the effect of an IGF-1-coding rAAV vector-laden alginate hydrogel (IGF-1/AlgPH155) in the treatment of full-thickness chondral defects [133]. Compared with the control (lacZ/AlgPH155) treatment, IGF-1/AlgPH155 consistently overexpressed IGF-1 for one year and thereby achieved an improvement in cartilage repair and a reduction in perifocal osteoarthritis and inflammation. In CTE, relatively little attention has been paid to the zonal organization of neocartilage tissue. To generate different types of neocartilage constructs, Weibenberger et al. fabricated a series of type I collagen hydrogels containing human bone marrow MSCs transfected with SOX9-, TGF-β1-, or BMP-2-encoding adenoviral vectors [134]. MSC chondrogenesis was detected in all hydrogels, but the expression of alkaline phosphatase (ALP) and Col-X demonstrated the three hydrogels had produced different types of neocartilage with varying levels of cartilage hypertrophy among them (BMP2 > TGFB1 > SOX9), indicating that this technology may be used for the regeneration of specific zones of native cartilage.

7.2. Nonviral Vector-Based Gene Delivery Platforms

The nonviral vector strategy requires the extraction of cells that were previously transfected with a gene vector, which are then introduced to the site of interest. Compared with viral-based gene delivery systems, nonviral systems are often considered to have advantages for in vivo use in cartilage regeneration because of their security, low immunogenicity, ease of operation, and relatively low cost [135, 136]. Nonviral gene therapies for CTE generally involve attaching chondrocytes or MSCs transfected with the desired gene for encoding chondrogenic growth factors onto hydrogel scaffolds. Madry et al. reported that the incorporation of human IGF-1 cDNA-transfected articular chondrocytes into alginate hydrogels enhanced the therapeutic effect of the hydrogel on full-thickness cartilage injuries in rabbits [137]. Lu et al. reported the use of gene therapy to achieve temporally controlled delivery of chondrogenic growth factors [138]. The authors developed a local gene delivery system using a porous chitosan scaffold embedded with hybrid HA/chitosan/pDNA nanoparticles that carry plasmids encoding TGF-β1. In vitro, the hydrogel sustained the release of the plasmids for more than 120 days. In addition, the development of chondrocytes and surrounding ECM structures was observed in the chitosan scaffolds. Gonzalez-Fernandez et al. performed gene delivery for osteochondral regeneration using TGF-β3 and BMP-2 in an MSC-loaded alginate hydrogel [139]. Their results showed that the genes were effectively delivered to the MSCs within the alginate hydrogels without any cytotoxicity. Furthermore, the delivery of TGF-β3 and BMP-2 significantly increased the sGAG and collagen production of the MSCs, especially when they were delivered in combination. The simultaneous regeneration of bone and cartilage is important for osteochondral defect repair but difficult to achieve in tissue regeneration [140]. Spatial delivery of genes using multiphase scaffolds for spatial guidance of cellular processes is a promising approach to facilitate the reconstruction of complex tissue structures. An effective strategy for the simultaneous regeneration of bone and cartilage using a gene delivery therapy was reported by Chen et al. [141]. The authors developed a bilayered scaffold consisting of a plasmid BMP-2-loaded hydroxyapatite/chitosan-gelatin scaffold for the osteogenic layer, and a plasmid TGF-β1-loaded chitosan-gelatin scaffold for the chondrogenic layer. MSCs in this bilayered scaffold were very proliferative, with TGF-β1 and BMP-2 proteins being highly expressed in their corresponding layer. Moreover, an osteochondral tissue composite was generated in vitro, and the osteochondral defect was repaired in vivo using this bilayered scaffold. This approach of spatially controlled gene delivery represents a promising strategy for osteochondral tissue engineering. An overview of the hydrogels used as gene delivery platforms for cartilage regeneration is summarized in Table 2.

8. Conclusions

Homeostasis and regeneration of cartilage tissue are strictly controlled by several signaling biomolecules and the interactions between cells and the ECM. The use of hydrogels as biomolecule delivery platforms is a promising strategy for improving the current treatments of CTE. This strategy can introduce appropriate levels of therapeutic biomolecules into target locations, and the specific dosage, spatial distribution, and temporal sequence can be precisely controlled. Moreover, hydrogel delivery platforms are highly tunable and can be designed to deliver biomolecules in specific ways by varying the hydrogel composition, crosslinking density, and degradation kinetics.

9. Future Perspectives

At present, researchers are gradually realizing that some biomolecules may only be suitable for certain delivery platforms, and the biocompatibility, activity, and stability of biomolecules are also key factors that determine the success of delivery platforms. Therefore, future efforts should be undertaken to identify the underlying mechanisms of biomolecule-to-platform interactions, thus improving the optimization of drug delivery platforms. Secondly, the degradation of hydrogels might impair their mechanical stability and long-term therapeutic effect; therefore, the incorporation of exogenous cells, such as MSCs, could potentially replace the degrading scaffolds with regenerated tissue. In addition, to successfully transfer hydrogel drug delivery systems from the laboratory to the market, further development is needed to design hydrogels with sufficient stiffness, minimal tissue response, and proper release and degradation profiles. Moreover, modular tissue assembly is a relatively unexplored strategy for stepwise biofabrication of 3D tissue constructs using cell-seeded scaffolds and preformed tissue units. The modular assembly enables the development of individually optimized and interchangeable modules that can jointly meet the requirements of functional cartilage regeneration scaffolds. Finally, noninvasive or minimally invasive diagnostic methods should be combined with the next generation of CTE, as the disease status and the therapeutic effect of the treatment could be observed and assessed in real time. Currently, the diagnosis of cartilage defects is mostly based on the results of imaging techniques, including X-ray, computerized tomography, magnetic resonance imaging, and ultrasound, but all of them have certain limitations. Thanks to the rapid development of artificial intelligence (AI) technology, its algorithm has been used to increase the accuracy of feature images. Moreover, the development of computational models and in silico simulations (such as the finite element method, computational fluid dynamics, or finite volume method) has enabled the rapid production of predictive models. It is reasonable to imagine that the combination of imaging techniques and these simulations may provide tissue engineers and physicians with timely, dynamic predictive models that can be used in the next generation of CTE.

Data Availability

No data were used to support the findings of the study.

Conflicts of Interest

The authors declare that they have no conflicts of interest.

Authors’ Contributions

Jingyan Guan and Jingwei Feng contributed equally to this work.


This work was supported by the National Natural Science Foundation of China (Grant nos. 81601702, 81671931, 81701920, 81801933, and 82102350).