The Functional Role of Interface Tissue Engineering in Annulus Fibrosus Repair: Bridging Mechanisms of Hydrogel Integration with Regenerative Outcomes

Abstract

Hydrogels are extraordinarily versatile by design and can enhance repair in diseased and injured musculoskeletal tissues. Biological fixation of these constructs is a significant determinant factor that is critical to the clinical success and functionality of regenerative technologies for musculoskeletal repair. In the context of an intervertebral disc (IVD) herniation, nucleus pulposus tissue protrudes through the ruptured annulus fibrosus (AF), consequentially impinging on spinal nerve roots and causing debilitating pain. Discectomy is the surgical standard of care to treat symptomatic herniation; however these procedures do not repair AF defects, and these lesions are a significant risk factor for recurrent herniation. Advances in tissue engineering utilize adhesive hydrogels as AF sealants; however these repair strategies have yet to progress beyond preclinical animal models because these biomaterials are often plagued by poor integration with AF tissue and lead to large variability in repair outcomes. These critical barriers to translation motivate this article to review the material composition of hydrogels that have been evaluated in situ for AF repair, proposed mechanisms of how these biomaterials interface with AF tissue, and their functional outcomes after treatment in order to inform the development of new hydrogels for AF repair. In this systematic review, we identify 18 hydrogel formulations evaluated for AF repair, all of which demonstrate large heterogeneity in their interfacing mechanisms and reported outcome measures to assess the effectiveness of repair. Hydrogels that covalently bond to AF tissue were found to be the most successful in improving IVD biomechanical properties from the injured state, but none were able to restore properties to the intact state suggesting that new repair strategies with innovative surface chemistries are an important future direction. We additionally review biomechanical evaluation methods and recommend standardization in the field of AF tissue engineering to establish mechanical benchmarks for translation and ensure clinical feasibility.

1. INTRODUCTION

Intervertebral disc (IVD) herniation is a highly prevalent and painful spine condition that affects approximately 20 in every 1000 adults per year.¹ Among the plethora of debilitating spinal pathologies that warrant surgical intervention, lumbar disc herniation remains the leading condition for which spine surgery is performed.² Discectomy is the current surgical treatment option to relieve low back pain and sciatica caused by herniation, for which over 400 000 procedures are performed annually in the United States alone.^3–5 In this procedure, a surgeon removes nucleus pulposus (NP) tissue fragments that protrude through fissures in the annulus fibrosus (AF), relieving mechanical compression of nerve roots in the spinal canal. Discectomy procedures are more effective in relieving neuropathy than nonoperative treatment, yet this procedure does not aim to repair defects in the AF.⁶ Unaddressed AF defects are a well-known contributor to undesirable postoperative outcomes including, but not limited to chronic low back pain, decrease in disc height, progressive disc degeneration, and recurrent herniation.^7–11 Of these complications, recurrent IVD herniation occurs most frequently in up to 23% of patients and is the leading cause of reoperation after discectomy.^12,13 Given the immense clinical significance and postoperative complications associated with AF defects, there is a clear gap in current surgical treatment methodologies that motivates the development of new treatment options for AF repair so as to improve patient outcomes.

Repairing the AF has historically been an intractable problem with no clinically viable solution on the market today, although there have been many previous and ongoing attempts to seal AF defects. The IVD’s anatomically complex structure, demanding mechanical loading behaviors, and avascular composition lead to a poor intrinsic healing capacity, where adult IVDs could neither recapitulate native tissue architecture nor restore mechanical function after injury.^14–16 This limited self-healing capacity necessitates reparative interventions that promote regeneration and restore organ function. Since the primary function of the IVD is mechanical, there have been a number of mechanical closure devices developed to repair annular lesions.^17–19 Namely, four mechanical devices with an indication for annular repair have been evaluated in preclinical or clinical settings for regulatory approval: X-Close Tissue Repair System, Inclose Surgical Mesh System, Disc Annular Repair Technology (DART) System, and the Barricaid Annular Closure Device (ACD).^20–31 The Inclose and X-Close devices received FDA 510(k) clearance in 2005 and 2009, respectively, and the Barricaid device was granted FDA premarket approval in 2019. To date, DART has not received FDA 510(k) clearance for marketing in the United States. When device effectiveness was assessed, these mechanical devices have demonstrated variable outcomes and limited success; in certain cases, these treatment options were not able to mitigate the risk of reherniation, restore functionally important biomechanical parameters such as disc height, or resist dorsal dislocation within the adjacent vertebrae upon initial ACD implantation.^27,29,32 Moreover, implantation of such devices involves a rather invasive surgical approach, requiring the surgeon to further damage AF tissue with suturing systems to affix the sutures to the outer radius of the AF or damage vertebral bodies to anchor titanium elements of the device. While these mechanical devices strictly aim to prevent recurrent herniation following discectomy, they fail to integrate with IVD tissue, recapitulate the native architecture of the IVD’s extracellular matrix (ECM), restore biomechanical behaviors to the IVD’s intact state, or promote healing. These shortcomings of biologically inert mechanical devices underscore the critical need to develop regenerative and minimally invasive therapies that can address the multitude of mechanical and biological requirements for successful AF repair.

The burgeoning field of tissue engineering and regenerative medicine led to the advent of adhesive hydrogel technologies for tissue repair, which hold specific advantages over mechanical devices as a next generation treatment option. They aim to provide structural support with an overarching goal to restore biological and biomechanical function to damaged tissues. As biomimetic water-swollen polymer networks, they can be mechanically tuned by their material composition for organ-specific applications and incorporate cellular and biologic factors to impart bioactivity and thereby enhance biological repair. These reparative biomaterials bear significant utility for surgical applications and are widely applied across numerous organ systems in addition to the IVD including, but not limited to, the eye, heart, liver, lungs, skin, and cartilage.^33–42 Traditional adhesives that draw upon natural and synthetic base materials such as fibrin glues, gelatin–resorcinol–formaldehyde/glutaraldehyde glues, cyanoacrylate glues, polysaccharide-based adhesives (gelatin-, collagen-, dextran-, and chitosan-derived, etc.), polypeptide-based adhesives, and polymeric adhesives (PEG-, acrylate-, and polyamine-based, etc.) have been investigated for surgical applications in these organ systems.^33,43–45 In addition to traditional adhesives, next generation materials are bioinspired by design and utilize tissue interfacing mechanisms that are naturally found in other animals such as mussels, slugs, spiders, geckos, barnacles, sandcastle worms, and caddisflies.⁴³

Adhesive hydrogels provide a distinct method of application to injured tissues as compared to mechanical devices and other polymeric scaffolds. These materials are often injectable, thereby lending themselves to a minimally invasive approach to repair soft tissues in the musculoskeletal system. This surgically convenient method of application renders them as a highly attractive option to repair focal AF defects following IVD herniation. Conversely, mechanical devices require invasive surgical methodologies to achieve durable fixation for AF repair, offering hydrogels a distinct advantage over these strategies provided that their performance is at least equivalent or superior to that of mechanical devices. Other AF repair strategies include electrospun or aligned fiber polymeric scaffolds, which are advantageous over hydrogels in that they can mimic the AF’s lamellar structure and anisotropic mechanical properties, yet they are not injectable or adherent and so not easily translatable to repair focal defects in the AF. These scaffolds are therefore better suited as an AF mimetic in tissue engineered composites for total disc replacement to treat late-stage IVD degeneration rather than AF defects arising from IVD herniation.^46–49

The traditional tissue engineering paradigm to develop hydrogel constructs in the context of AF repair is shown in Figure 1, where hydrogels are composed of prepolymer solution containing natural or synthetic polymers that can embed cells or biochemical factors within their three-dimensional matrixes with the potential addition of biomechanical stimuli. These engineered constructs for AF repair can be tailored for biomechanical restoration or biologic delivery, where the biomechanical design requirements for AF repair have been extensively described and are distinctly different than those for the successful delivery of cells or other biologics.^18,50 However, avoiding neuropathy from herniation of IVD tissues or biomaterial implants is of paramount importance. Therefore, in order for these biomaterials to have any appreciable clinical utility, at minimum, they need to be retained within the site of the AF defect under the distinct physiological range of axial loads, torsional loads, and motion segment bending.^50–53 This essential requirement for translation underscores the specificity for AF tissue engineering design criteria as reported by Long et al., which are distinctly different than those used to develop surgical adhesives for other tissue systems (i.e., eye, skin, heart, lungs, liver, etc.).⁵⁰

Figure 1.

Tissue engineering paradigm that guides the fabrication of regenerative constructs for AF repair. Image used with permission from Mount Sinai Health System.

Hydrogels must durably adhere to AF tissue surfaces in order to meet this required benchmark for translation and not dislodge from the repair site after delivery. Although an engineered construct may demonstrate suitable mechanical or biological performance in vitro, the biomaterial must integrate with the native fibrocartilage so that the likelihood of implant herniation, and thus reoperation, is minimized. Hydrogel adherence to wetted tissue surfaces is imparted by physical or chemical interactions across a range of length scales, where constitutive polymer network(s) feature intrinsic adhesivity based upon molecular composition and can be engineered to optimize these adherent interactions. When we consider the complex architecture of the IVD, hydrogels have to not only serve as an adhesive biomaterial but also function as a sealant so as to prevent reherniation. Often these terms are used interchangeably; however these concepts are functionally distinct from one another. By definition, a sealant establishes a physical barrier to prevent leakage of a given substance, whereas an adhesive conjoins two similar or dissimilar surfaces.⁴³ Given the IVD’s anatomical structure and biomechanical behavior upon loading, these functional requirements are inextricably related, and the hydrogel must function as both a tissue sealant and adhesive when applied to AF defects. Upon axial compression, NP pressurization induces radially outward forces that are transferred to the AF, where the AF consequentially bears internal loads in circumferential tension along the direction of its lamellae. Therefore, in order to successfully repair the AF, hydrogels must retain NP tissue centrally within the IVD (ability to seal AF defects) as well as bridge the surfaces of AF tissue together within the interior of the defect space (ability to adhere to AF tissue), in which these essential functions are contingent on interfacial properties between the hydrogel and AF as well as the hydrogel’s bulk material properties. Thus, biomechanical evaluation of the disc upon physiological loading as well as additional assessments of implant herniation risk and durability is of major importance to test the capacity of new hydrogel formulations to properly seal AF defects and adhere to AF tissue.^50,52

Although the material composition and interfacing mechanisms for the aforementioned adhesives have been described, AF repair hydrogels represent a distinct class of these biomaterials and only a fraction of those investigated for AF repair are reported in the literature. To date, there has been no comprehensive overview in the field of interface tissue engineering that reports (1) the distinct polymeric compositions for AF repair hydrogels that have been evaluated in a preclinical setting, (2) their interfacing mechanisms with AF tissue, and (3) their functional outcomes after application to AF defects. Here, we address these gaps in the literature by providing a systematic review that details the polymeric composition of hydrogels previously evaluated for AF closure, reports associated functional outcomes for this set of biomaterials, and provides biomechanical testing methodologies to advance these tissue engineering strategies toward clinical application. Additionally, we review the physical and chemical bases of adhesion theory in tandem with the proposed interfacing mechanisms and surface chemistries of hydrogels investigated for AF repair with the goal to inform the design and application of new bioadhesive polymer networks.

2. LITERATURE REVIEW METHODS

PubMed, Scopus, and MEDLINE were the three literature databases used in this systematic review, from which a total of 3554 citations were identified from the primary search in July 2019 (and subsequently updated in December 2019 and October 2020) using the following search terms: (1) “Annulus Fibrosus” AND “Repair”, (2) “Annulus” AND “Hydrogel”, (3) “Annulus” AND “Plug”, and (4) “Intervertebral Disc” AND “Repair” (Figure 2). After careful examination of each citation according to the inclusion and exclusion criteria, 38 non-duplicate original research articles were retrieved corresponding to a total of 18 hydrogels evaluated for AF repair.

Figure 2.

Literature review methods for this systematic review article.

3. BIOMATERIAL INTEGRATION: STRUCTURE DICTATES ADHESIVE FUNCTION

The ability of a hydrogel to function as a surgical sealant and adhesive is contingent on the extent of integration with the native tissue, where integration is formally defined as the connectivity between two materials of homogeneous or disparate composition.⁵⁴ Integration is conceptually focused on the biological interface between the tissue in need of repair and an implantable hydrogel system. Although great effort has been spent on the development of hydrogel systems as viable substrates for tissue engineering applications in vitro, far less scientific attention has been directed at the interface between these hydrogel systems and tissue of interest. Yet, the potential for clinical translation is dependent on establishing a durable connection at the tissue–biomaterial interface so as to ensure construct longevity and satisfactory postoperative performance. As tissue engineering and regenerative medicine strategies advance toward clinical applications in orthopedics, investigators should place emphasis on the role of biomaterial–tissue interfaces to ensure translational success. Given the musculoskeletal system’s primary roles in providing structural support, distributing internal loads, and enabling locomotion, it is critical to form and maintain a robust interface between the biomaterial and tissue surface. Successful integration should render sturdy biological fixation that manifests in hydrogel immobilization in situ, thereby enabling immediate tissue functionality upon implantation.⁴¹

When designing hydrogel systems for integration, one must consider the possible mechanisms by which tissue integration can be accomplished. Integration is functionally imparted by biomaterial adhesion, where the adhesion of hydrogel substrates onto tissue surfaces can be achieved by either physical or chemical means spanning multiple length scales. Here, we present the three modalities of traditional adhesion theory: mechanical interlocking, electrostatic interactions, and chemical interactions⁵⁵ (Figure 3). The goal of providing this scientific overview of adhesion theory is to inform the selection and engineering of polymeric biomaterials for next generation AF repair hydrogels, which may utilize any combination of these biophysical and biochemical means of adhesion.

Figure 3.

Classical adhesion theory and associated chemical bonds that can arise from surface chemistries at the tissue–hydrogel interface, with chemical bonding mechanisms partly informed by Ghobril and Grinstaff.⁵⁶ Image used with permission from Mount Sinai Health System.

Mechanical interlocking and electrostatic interactions arise from biophysical phenomena at the interface between the adhesive hydrogel and tissue surface.^56,57 Mechanical interlocking occurs on a microscopic scale of action, where surface roughness of the adherend permits interdigitation of the hydrogel within surface irregularities and imparts adhesion to tissue surfaces. Experimental methods to visualize mechanical interlocking between the adhesive hydrogel and tissue include optical or electron microscopy. Electrostatic interactions occur on the molecular scale of action, where atomic differences in electronegativities establish an asymmetric distribution of electrons on adhesive and adherend surfaces. This electrostatic asymmetry results in partial positive and partial negative charges between surfaces that attract one another and functionally resist separation.⁵⁷

Chemical interactions are the third broad class of mechanisms pertaining to adhesion theory and arise from biochemical phenomena between the hydrogel adhesive and tissue surface at the interface. These chemical interactions can be characterized by diffusion, physisorption, and chemisorption, which occur on atomic and molecular scales of action.⁵⁸ Diffusion arises primarily between polymeric adhesives and adherends, where the constitutive polymers of the hydrogel network and biopolymers of the tissue interpenetrate one another at the interface (i.e., interdiffusion).⁵⁹ Physisorption corresponds to the physical basis of adsorption, where noncovalent intermolecular interactions impart biomaterial adhesion onto a tissue substrate.⁶⁰ These noncovalent intermolecular forces can arise from hydrogen bonding or van der Waals forces at the interface.⁶¹ Chemisorption corresponds to the chemical basis of adsorption, where covalent linkages impart hydrogel adhesion onto tissue substrates.⁶⁰ There are a number of surface chemistries that can occur between the adhesive hydrogel and tissue, which can result in many types of covalent bonds, including imine bonds, amide bonds, urea bonds, N–N bonds resulting in hydrazine derivatives, bonds arising from Michael addition, and disulfide bridges arising from thiol oxidation.⁵⁶ These covalent bonds can form by reacting specific functional groups in the hydrogel network with either primary amines or thiol groups found on extracellular matrix proteins at the tissue surface.

The primary mechanism(s) by which a hydrogel adheres onto AF tissue is prescribed by the biochemical structure, biophysical conformation, and molecular composition of its polymer network(s) in addition to the surface topography of the AF. Hydrogels screened in situ for AF repair vary in composition, and constitutive networks can be composed of natural polymers, synthetic polymers, or a combination of the two polymer types. Since the polymer repeat units and functional groups are specific to the macromers of a given hydrogel formulation, each hydrogel has unique physical and chemical interactions with the AF at the tissue surface rendering different mechanisms of adhesion across formulations. These differences in integration mechanisms are most likely to demonstrate differences in biomechanical outcomes following AF repair due to the strong association between integration mechanism and hydrogel adhesivity. Biomechanical parameters that quantify construct herniation risk would be the most relevant output measures to distinguish integration mechanisms with respect to their functional efficacy for repair. The composite list of hydrogels screened in situ for AF repair and their proposed mechanisms of adhesion are presented in Table 1. Gross specimen visualization and histological analysis of each repair strategy depict the large heterogeneity in composition and adhesiveness across hydrogel formulations that have been developed for AF repair (Figure 4).

Table 1.

Hydrogels Evaluated in Situ for AF Repair and Proposed Mechanisms of Adhesion

Hydrogel composition	Abbreviation	Primary mechanism(s) of adhesion	Test method of application	Species of IVD model	Refs
genipin-cross-linked fibrin	FibGen	chemisorption	ex vivo and in vivo	bovine, ovine	62–73
riboflavin-cross-linked collagen	RF-collagen	chemisorption	ex vivo and in vivo	rat, ovine	74–82
Rose Bengal-cross-linked collagen	RB-collagen	chemisorption	ex vivo and in vivo	rabbit	83
poly(ethylene glycol)–poly(trimethylene carbonate)–hexamethylene diisocyanate	PEG400-TMC3-HDI	chemisorption	ex vivo	bovine, caprine	84,85
albumin/glutaraldehyde	BioGlue	chemisorption	ex vivo	bovine	69
n-butyl cyanoacrylate (n-BCA) and octyl cyanoacrylate (OCA)	cyanoacrylate/LiquiBand	chemisorption	ex vivo and in vivo	porcine	86
citric acid–1-ethyl-3-(3-dimethylaminopropyl) carbodiimide/N-hydroxysuccinimide type I collagen	CA-EDC/NHS collagen	chemisorption	in vivo	rat	87
1-(3-dimethylaminopropyl)-3-ethyl-carbodiimide hydrochloride-cross-linked gelatin/poly(γ-glutamic acid)	EDC-gelatin/γPGA	chemisorption	ex vivo	bovine	88
fibrin	fibrin	chemisorption	ex vivo and in vivo	porcine, bovine, ovine	89,90
poly(lactic-co-glycolic acid)/fibrin	PLGA/fibrin	chemisorption	in vivo	rabbit	91
poly(ethylene glycol) diacrylate + chondroitin sulfate or hyaluronic acid methacrylate aldehyde	PEGDA/fibrin + CSMA or HAMA aldehyde	chemisorption	ex vivo	bovine	92
poly(lactic-co-glycolic acid)	PLGA	electrostatic interactions, physisorption, and diffusion	in vivo	rabbit	93
alginate/collagen	alginate–collagen	electrostatic interactions, physisorption	ex vivo	bovine	94
alginate	alginate	electrostatic interactions, physisorption	ex vivo	bovine	95
hyaluronic acid	HA	electrostatic interactions, physisorption, and diffusion	ex vivo and in vivo	bovine, rat	96,97
hyaluronic acid/albumin	HA/albumin	electrostatic interactions, physisorption, and diffusion	in vivo	ovine	98
ultrapurified alginate	UPAL	electrostatic interactions, physisorption, and diffusion	ex vivo and in vivo	rabbit	99
cellulose nanofiber-reinforced chitosan	CNF/CHI	electrostatic interactions, physisorption	ex vivo	porcine	100

Figure 4.

Gross specimen and histological images of repaired IVDs that were treated with a given hydrogel formulation. Reproduced with permission from ref 66, Copyright 2014 the Authors; ref 69, Copyright 2018 the Authors; ref 76, Copyright 2017 Elsevier; ref 79, Copyright 2015 Elsevier; ref 83, Copyright 2013 Elsevier; ref 84, Copyright 2018 John Wiley & Sons, Inc.; ref 86, Copyright 2017 John Wiley & Sons, Inc.; ref 87, Copyright 2017 Springer Nature; ref 90, Copyright 2011 Wolter Kluwer Health, Inc.; ref 91, Copyright 2017 Springer Nature; ref 92, Copyright 2020 Elsevier; ref 93, Copyright 2016 Springer Nature; ref 94, Copyright 2015 SAGE; ref 95, Copyright 2014 Elsevier; ref 99, Copyright 2018 Elsevier; ref 100, Copyright 2019 MDPI.

4. FUNCTIONAL EVALUATION METHODS

Following their molecular design and formulation, candidate hydrogels are first evaluated in vitro and subsequently assessed in situ using ex vivo or in vivo preclinical animal models. Guidance documentation published by international standards organizations and regulatory agencies aims to provide a systematic framework for investigators to advance new tissue engineering technologies along the translational development process by establishing functional evaluation methods. Although strict adherence to guidance documentation is not required for regulatory approval, it aids in the development of robust study designs and the selection of relevant testing configurations and specifications as well as experimental output measures to facilitate efficient regulatory review. Without this form of documentation, investigators follow individual or lab-specific preferences, resulting in discordance across study outcomes and making it difficult to compare the effectiveness of treatment across hydrogels of different material compositions.¹⁰¹ Since this type of documentation does not yet exist to guide the development and evaluation of new hydrogel systems specific for AF repair, we provide in vitro and in situ biomechanical evaluation methods to establish best practices when screening the performance of newly developed hydrogel formulations and to ensure that these materials meet minimum feasibility benchmarks for translation. By providing this information, we aim to promote congruence in study outcomes and facilitate comparison among future studies in the field of AF tissue engineering.

4.1. Direct Assessment of Hydrogel Adhesion to Tissue.

Integration can be functionally assessed with mechanical tests that determine adhesion strength properties between adhesive hydrogels and tissue adherends. ASTM International established five standard specifications of different testing configurations to evaluate interfacial strength properties for tissue adhesives and surgical sealants. Investigators should use these standards as an initial assessment of tissue–hydrogel adhesivity and as a screening tool to evaluate hydrogel formulations of interest before translating these biomaterials into a preclinical animal model. Although all ASTM standards provide a measure of hydrogel adhesivity to tissue substrates, certain ASTM tests could be more relevant as a screening tool for AF repair hydrogels. Since NP pressurization induces a radial outward force on the implanted hydrogel, ASTM 2255–05 and ASTM 2392–04 are the most similar in vitro tests that recapitulate this type of biomechanical loading in situ, given the direction of applied loads with respect to the adhesive hydrogel in these two testing configurations.

4.1.1. Single and Double Lap Shear, ASTM F2255–05.

The Lap Shear adhesion test corresponds to ASTM F2255–05, which delineates the testing specifications and guidelines for the “Strength Properties of Tissue Adhesives in Lap-Shear by Tension Loading.”¹⁰² The lap shear test is useful for mechanical characterization of adhesives. The ASTM F2255–05 standard is designed for adhesives meant to be used as surgical sealants on soft tissues. At least ten specimens of each type are to be tested. The length of the tissue substrate attached to each specimen holder should be 1.5 times the length of the bond area (1.0 ± 0.1 cm), the tissue specimens should be 1–2 mm thick, and tissue substrate should be attached to the specimen holder with a suitable adhesive. A bond force of 1–2 N is placed on the bond area between the two tissue specimens (1.0 ± 0.1 cm by 2.5 ± 0.1 cm) until the experimental adhesive sets. The specimens are conditioned for 1 h ± 15 min in phosphate buffered saline at 37 °C ± 1 °C. After conditioning, samples are acclimated to the test temperature for 15 min. The sample is loaded into the testing apparatus such that the load coincides with the long axis of the sample. The sample is loaded to failure at a constant cross-head speed of 5 mm/min. The load at failure (maximum load sustained) and the type of failure (percentage cohesive, adhesive, or substrate failure) are recorded (Figure 5A).

Figure 5.

Biomechanical evaluation methods for adhesive hydrogels. (A–E) In vitro ASTM standards to directly assess hydrogel adhesion. (F) Testing configuration to evaluate the herniation risk of hydrogel repair strategies. (G) Testing configuration to evaluate the durability of hydrogel repair strategies. Image used with permission from Mount Sinai Health System.

4.1.2. Butt Joint Test: Pull Off Resistance, ASTM F2258–05.

The Butt Joint adhesion test corresponds to ASTM F2258–05, which describes testing specifications and guidelines for the “Strength Properties of Tissue Adhesives in Tension.”¹⁰³ This test is useful for mechanical characterization of adhesives. At least ten specimens of each type are to be tested. The only critical dimension in this type of test is the bond area of 2.5 ± 0.005 cm by 2.5 ± 0.005 cm. A bond force of 1–2 N is applied until the experimental adhesive sets. The specimens are conditioned for 1 h ± 15 min in phosphate buffered saline at 37 ± 1 °C. After conditioning, samples are acclimated to the test temperature for 15 min. The sample apparatus is loaded into the tensile test machine and at a constant cross-head speed of 2 mm/min. The load at failure (maximum load sustained) and the type of failure (percentage cohesive, adhesive, or substrate failure) are recorded (Figure 5B).

4.1.3. T-Peel Joint Test, ASTM F2256–05.

The T-Peel adhesion test corresponds to ASTM F2256–05, which delineates the testing specifications and guidelines for the “Strength Properties of Tissue Adhesives in T-Peel by Tension Loading.”¹⁰⁴ This test is useful for mechanical characterization of adhesives. At least ten specimens of each type are to be tested. Tissue specimen thickness should be uniform and less than 5 mm. The specimen width is 2.5 ± 0.1 cm, and the specimen length is 15 ± 0.2 cm (2.5 cm unbonded, 12.5 cm bonded). A bond force of 5–10 N is applied until the experimental adhesive sets. The specimens are conditioned for 1 h ± 15 min in phosphate buffered saline at 37 ± 1 °C. After conditioning, samples are acclimated to the test temperature for 15 min. The sample apparatus is loaded into the tensile test machine and at a constant cross-head speed of 250 mm/min. The load as a function of displacement and the type of failure (percentage cohesive, adhesive, or substrate failure) are recorded (Figure 5C).

4.1.4. Wound Closure Test, ASTM F2458–05.

The wound closure test corresponds to ASTM F2458–05, which describes testing specifications and guidelines for the “Standard Test Method for Wound Closure Strength of Tissue Adhesives and Sealants.”¹⁰⁵ This test is useful for assessing the wound closure strength of tissue adhesives used to help secure the apposition of soft tissue. At least ten specimens of each type are to be tested. Two tissue samples of identical size (10 ± 0.2 cm by 2.5 ± 0.1 cm) are bonded using the experimental adhesive on the 2.5 cm side, with a bonding length of 0.5 cm on either side of the join line. The thickness of the specimens should be uniform and less than 5 mm. The specimens are conditioned for 1 h ± 15 min in phosphate buffered saline at 37 ± 1 °C. After conditioning, samples are acclimated to the test temperature for 15 min. The sample is loaded into the testing apparatus such that the load coincides with the long axis of the sample. The distance from the grip to the midline of each sample is 5 cm, with the remaining 5 cm held between the grips. The specimen is loaded to failure at a constant speed of 50 mm/min. The time from application to testing (cure time), force at failure (maximum force required to disrupt substrate), and the type of failure (percentage cohesive, adhesive, or substrate failure) are recorded (Figure 5D).

4.1.5. Burst Strength Test, ASTM F2392–04.

The Burst Strength test corresponds to ASTM F2392–04, which describes testing specifications and guidelines for the “Standard Test Method for Burst Strength of Surgical Sealants.”¹⁰⁶ This test is used to evaluate the burst or rupture strength of sealants on soft tissue. At least ten specimens of each type are to be tested. This test employs an apparatus that clamps down on a substrate to prevent leakage. The thickness of the tissue should be uniform and not exceed 5 mm. Tissue samples should be circles 3.0 ± 0.1 cm in diameter, in which a 3.0 mm diameter hole is created using a biopsy punch. The specimens are conditioned for 1 h ± 15 min in phosphate buffered saline at 37 ± 1 °C. After conditioning, samples are acclimated to the test temperature for 15 min. This test uses a stationary fixture containing test substrate and the sealant to be tested. A 1.0 mm thick PTFE mask with a 15 mm diameter hole is secured over the sample, with the hole in the mask centered with the hole in the sample. Saline is pumped into the fixture at a constant rate of 2 mL/min, and pressures are measured at all time points. Peak pressure and failure type (cohesive, adhesive, or substrate) are recorded (Figure 5E).

4.2. Assessment of Implant Herniation Risk and Durability.

Although the aforementioned ASTM standards are useful to directly assess the adhesion strength properties of candidate hydrogel formulations, they are contrived in vitro testing configurations that do not evaluate performance in a physiologically relevant in situ context for AF repair. To translate candidate hydrogels into a preclinical animal model, it is critically important to determine in situ AF repair, a configuration that does not currently have ASTM standards, by assessing the following functional outcomes: technical feasibility, implant herniation risk, and construct durability.

4.2.1. Herniation Risk: Ramp-to-Failure Testing.

IVD herniation risk following AF repair is assessed by a displacement-controlled ramp-to-failure test first reported by Vergroesen and colleagues⁸⁵ and subsequently implemented by other groups.^62,84,107 In this ex vivo test, the motion segment undergoes compression at a fixed rate of 2 mm/min until the ultimate strength of the specimen is reached. The motion segment is compressed under 5° left lateral flexion, where the NP exerts an appositional force against the hydrogel repair site (Figure 5F). This biomechanical test is designed to assess the failure properties of the IVD motion segment and sealing capacity of the adhesive hydrogel in the worst-case scenario because motion segment bending directs NP displacement in the radial direction of the hydrogel. Biomechanical output measures that quantitatively describe IVD herniation risk include failure strength, failure strain, subsidence-to-failure, maximum stiffness, work-to-failure, yield strength, ultimate strength, and the ratios of the ultimate or yield strength to the failure strength of the motion segment. When conducting this test, investigators should compare repaired IVD motion segments against intact IVDs as a control to evaluate the extent of biomechanical restoration and additionally compared against an IVD injury model simulating discectomy to demonstrate biomechanical noninferiority or superiority to the surgical standard of care.

4.2.2. Durability: Fatigue Endurance Testing.

The longevity of the AF repair strategy is evaluated by fatigue endurance testing, where the IVD motion segment undergoes cyclic bending until failure is reached. The biomechanical output measure to evaluate repair durability is the number of cycles-to-IVD failure. Wilke and colleagues first established a reliable in vitro herniation model and evaluated durability of the Barricaid Anular Closure Device. In this test, informally referred to as the “hula hoop test”, cyclic loading is applied to the IVD motion segment to a total of 100 000 cycles and paused at 20 000, 40 000, 60 000, and 100 000 loading cycles to inspect for the NP extrusion, conduct flexibility tests, and measure disc height.¹⁰⁸ Cyclic testing is performed in a servohydraulic material testing machine, where motion segment specimens are fixed with flanges in a rotating base that is mounted transversely and rotates at 360 deg/min. The rotation base is shifted 30 mm laterally to apply eccentric loads to the motion segment. The hydraulic piston first linearly ramps force up to a preload of 100 N and then applies a sinusoidal force that ranges between 100 and 600 N at 5 Hz. The 30 mm shift in applied force resulted in applied bending moments that reached a maximum of 18 N·m. An adapted protocol by Lin and colleagues,¹⁰⁷ for systems unable to apply continuous rotation, applies cyclic eccentric compression with a loading indenter that rotates ±135° from the axis opposite of the defect site at 15° increments with 1 min of cyclic loading at each location. Magnitudes of cyclic compression range between 50 and 300 N at 1 Hz at a 20 mm offset distance from the center of the IVD to induce a physiologically relevant bending moment of 6 N·m (Figure 5G).¹⁰⁷ Similar to herniation risk assessments, evaluation of fatigue endurance should compare repaired IVD motion segments against intact IVD controls as well as unrepaired IVD motion segments that simulate discectomy to demonstrate biomechanical noninferiority or superiority to the surgical standard of care.

5. BIOCHEMICAL MECHANISMS OF HYDROGEL INTEGRATION WITH THE AF

The molecular composition of each hydrogel system prescribes how the biomaterial will chemically interface with AF tissue due to differences in functional groups and repeat units on the polymer backbone. In this section, we propose the chemical mechanisms of hydrogel adsorption to AF tissue (i.e., chemisorption) for all evaluated biomaterials that result in covalent bonding with AF matrix proteins. For simplicity, we represent reactive primary amines on collagen since it is the most abundant protein in the AF; however these reactions can take place with other proteins present in the extracellular matrix of the AF.

5.1. Genipin-Cross-Linked Fibrin (FibGen) Hydrogels.

Iatridis and colleagues developed genipin-cross-linked fibrin (i.e., FibGen) hydrogels for AF repair, after it was previously described for cartilage tissue engineering by Dare et al., and this biomaterial was additionally investigated for AF repair by Scheibler et al. and Frauchiger et al.^62–73,109 FibGen is composed of natural biopolymers (fibrinogen) and natural compound derivatives (genipin). After injecting the prepolymer solution into the AF defect, the fibrinogen monomers undergo thrombin-mediated polymerization to form the primary fibrin network, and this natural polymer network chemically binds to AF tissue by means of transglutaminase 2 activity as described in section 5.9. Genipin molecules simultaneously cross-link the hydrogel’s backbone by reacting with primary amines on fibrin and subsequently dimerizing. In addition to hydrogel cross-linking, the biomaterial is also covalently bound to AF tissue by genipin dimerization through the same means. Butler et al. proposed mechanisms of genipin-mediated cross-linking, where primary amines on either AF collagen or fibrinogen macromer perform a nucleophilic attack on the C3 carbon in the dihydropyran ring and open the heterocyclic molecule to generate an intermediate aldehyde.¹¹⁰ The secondary amine in the intermediate complex then attacks the aldehyde moiety forming a new heterocyclic compound, where collagen protein or fibrinogen macromers are then conjugated to the genipin molecule. In addition to thrombin-mediated polymerization, the fibrin gel is further stabilized when fibrin-bound genipin molecules dimerize, as described by the diffusion reaction model.¹¹¹ FibGen hydrogels adhere onto AF tissue when fibrin-bound genipin molecules dimerize with collagen-bound genipin molecules at the tissue–biomaterial interface. Since the extent of covalent bridging between the AF and hydrogel is dependent genipin dimerization, it logically follows that biomaterial adhesion is proportional to the genipin concentration in the hydrogel. This is corroborated by biomaterial push-out strength experiments, where an increase in FibGen push-out strength was observed with an increase in genipin concentration.⁶⁸ This is also supported by implant herniation risk experiments, where a significant increase in IVD failure strength was observed for IVD specimens repaired with FibGen of higher genipin concentrations (6 mg/mL) compared to lower genipin concentrations (1 mg/mL), while holding the fibrin concentration constant (Figure 6).⁶²

Figure 6.

Integration mechanism for FibGen hydrogels.

5.2. Riboflavin-Cross-Linked Collagen Hydrogels.

Bonnassar and colleagues developed an injectable riboflavin-cross-linked rat-tail type I collagen hydrogel for AF repair.^74–82 The rat-tail type I collagen macromers are covalently cross-linked to AF collagen and other collagen macromers through three possible photochemical mechanisms in the presence of riboflavin cross-linker (Figure 7).

Figure 7.

Integration mechanism for RF-cross-linked collagen hydrogels.

5.2.1. Pathway 1.

Upon exposure to singlet-state oxygen and long-wavelength ultraviolet radiation (UVA), riboflavin (RF) yields a self-activated photoproduct, 2,3-butanedione, that serves as a dielectrophile for subsequent reactions with collagen amino acid residues.¹¹² The primary structure of AF collagen and rat-tail type I collagen contains allysine and hydroxyallysine residues that result from endogenous lysyl oxidase activity and undergo aldol addition with 2,3-butanedione followed by an aldol condensation reaction. The symmetric nature of 2,3-butanedione lends itself to two aldol addition and condensation reactions thereby forming a covalent cross-link between the hydrogel and AF after UVA exposure.

5.2.2. Pathway 2.

Prior studies suggest that the imidazole group on the histidine side chain is highly amenable to reactions with singlet-state oxygen, evidenced by a significant reduction of histidine content in the presence of UVA-irradiated collagen and RF.^113,114 Upon reaction, the imidazole ring is oxidized and transformed into an imidazolone moiety, which then serves as an electrophile to react with nucleophilic amino acids. Residues that likely perform a nucleophilic attack on the imidazolone ring contain secondary or phenolic alcohols such as hydroxyproline, threonine, and tyrosine, among others. This nucleophilic attack results in a covalent cross-link between AF and hydrogel collagen structures, partly responsible for a demonstrated increase in stromal tensile strength of corneal tissue after exposure to UVA in the presence of RF and singlet oxygen.¹¹²

5.2.3. Pathway 3.

This pathway is analogous to pathway 2 in that the intermediate imidazolone moiety undergoes a nucleophilic attack with allysine or hydroxyallysine residues after lysyl oxidase yields its lysine derivates. This covalent linkage between collagen in the AF and hydrogel is identical to that of pathway 2.

5.3. Rose Bengal-Cross-Linked Collagen Hydrogels.

Chan and colleagues developed a Rose Bengal-cross-linked collagen hydrogel for cell delivery applications to the AF composed of rat-tail type I collagen.⁸³ Rose Bengal (RB) is a photosensitive cross-linking reagent that bonds collagen proteins together after exposure to green light (RGX). RGX excitation creates an intermediate singlet excited state RB (¹RB⁻²) that quickly undergoes an intersystem crossing to become triplet excited state RB (³RB⁻²). ³RB⁻² functionally serves as both an oxidizing and a reducing reagent for redox reactions with collagen’s amino acid residues, which produce free radical forms of amino acids. Oxidized and reduced amino acid radicals then undergo deprotonation and protonation, respectively, which leads to the termination step to create collagen cross-links at the AF-hydrogel interface or within the hydrogel construct itself.¹¹⁵ (Figure 8) Notably, the Rose Bengal cross-linking mechanism is far less understood than other photochemical strategies to cross-link collagen (e.g., riboflavin-based cross-linking).

Figure 8.

Integration mechanism for RB-cross-linked collagen hydrogels.

5.4. Poly(ethylene glycol) and Poly(trimethylene carbonate) Copolymer Hydrogels.

Smit and colleagues developed a copolymer hydrogel system for AF closure, and this formulation was further investigated by Long et al. for repair endurance, herniation risk, cytocompatibility, and biomechanical restoration.^84,85 The copolymer system (PEG₄₀₀-TMC2/3-HDI) is composed of poly(ethylene glycol) [PEG₄₀₀], poly-(trimethylene carbonate) [TMC2/3], and a hexamethylene diisocyanate (HDI) end group. Upon injection to AF defects, macromers undergo carbamate formation to form urethane bonds that cross-link macromers and induce gelation in situ. At the AF–hydrogel interface, HDI end groups undergo carbamide formation, which creates a covalent linkage between AF collagen and the hydrogel network through urea bonds (Figure 9).

Figure 9.

Integration mechanism for PEG₄₀₀-TMC3-HDI hydrogels.

5.5. Albumin–Glutaraldehyde (BioGlue) Hydrogels.

Scheibler et al. performed annular repairs with BioGlue (a commercially available sealant composed of albumin–glutaraldehyde) and compared biomechanical repair performance to FibGen-repaired specimens.⁶⁹ BioGlue components are applied to tissue defects through a dual-barrel syringe, which homogeneously mixes bovine serum albumin (BSA) and glutaraldehyde to initiate carrier conjugation. In step 1, end-group aldehyde moieties from glutaraldehyde undergo Schiff base formation with primary amines from BSA to form a two-part globular complex that will (1) further undergo Schiff base formation with adjacent albumin–glutaraldehyde complexes to form an adhesive hydrogel network and (2) form imine bonds upon reacting with primary amines on extracellular matrix proteins (i.e., collagen, elastin, etc.) in the AF (step 2) (Figure 10).

Figure 10.

Integration mechanism for albumin–glutaraldehyde (BioGlue) hydrogels.

5.6. Cyanoacrylate-Based (LiquiBand) Adhesive.

Kang and colleagues investigated the use of LiquiBand for AF repair, which is a cyanoacrylate-based adhesive composed of n-butyl cyanoacrylate (n-BCA) and 2-octyl cyanoacrylate (2-OCA).⁸⁶ The mechanism of action follows the standard scheme of polymerization, where adsorption and network formation is achieved through initiation, propagation, and termination. When applied to the AF surface, primary amines and thiol groups can initiate the polymerization process and undergo an aza-Michael addition or thiol-Michael addition reaction, respectively, resulting in an unstable nucleophile at the β-carbon position and inducing propagation. This reactive nucleophile then attacks unreacted n-BCA and 2-OCA monomers, thus creating new carbon–carbon bonds and establishing a polymerized network within a relatively short time frame. Following propagation, hydronium ions then terminate the polymerization reaction, where the unstable β-carbon nucleophile is protonated resulting in a stable and highly cross-linked network between n-BCA, 2-OCA, and AF collagen at the tissue surface (Figure 11).

Figure 11.

Integration mechanism for cyanoacrylate (LiquiBand) hydrogels.

5.7. CA-EDC/NHS Collagen Hydrogels.

Wang and colleagues investigated the use of rat-tail type I collagen hydrogels combined with citric acid (CA), 1-(3-dimethylaminopropyl)-3-ethyl-carbodiimide hydrochloride (EDC), and N-hydroxysuccinimide (NHS) for AF repair in an in vivo rat-tail model.⁸⁷ In this study, they assessed hydrogel formulations with and without the incorporation of CA for downstream experimentation. For formulations that include CA, EDC reacts with CA to form an unstable disubstituted or trisubstituted o-acylisourea intermediate. In the case of disubstituted CA, the intermediate reacts with NHS to form a semistable amine-reactive NHS-ester. In the presence of two macromers of rat-tail type I collagen, the dual-NHS-ester intermediate reacts with primary amines to form a covalent cross-link between macromers and polymerize the hydrogel network. At the tissue–hydrogel interface, the dual-NHS-ester intermediate reacts with primary amines on both rat-tail type I collagen macromers and primary amines found on extracellular matrix proteins in the AF, forming a stable amide bond between the hydrogel and AF. In the case of trisubstituted CA, the unstable intermediate similarly reacts with NHS to form a semistable amine-reactive NHS-ester at three positions. At the tissue–hydrogel interface, this configuration lends itself to two possible products that bridge two macromers of rat-tail type I collagen with one molecule of AF collagen or one macromer of rat-tail type I collagen with two molecules of AF collagen. For formulations that do not incorporate CA, EDC first reacts with the C-terminus, aspartic acid, and glutamic acid residues on rat-tail type I collagen and AF collagen to produce the unstable o-acylisourea intermediate. This intermediate product reacts with NHS to form the semistable amine-reactive NHS-ester, which then reacts with primary amines on rat-tail type I collagen to form a polymerized hydrogel network or primary amines on AF collagen to form a covalent bridge at the tissue–hydrogel interface (Figure 12).

Figure 12.

Integration mechanism for CA-EDC/NHS-cross-linked collagen hydrogels.

5.8. EDC-Gelatin-γPGA Hydrogels.

Yang and colleagues developed a 1-(3-dimethylaminopropyl)-3-ethyl-carbodiimide hydrochloride (EDC)-cross-linked gelatin and poly(γ-glutamic acid) (gelatin-γPGA) hydrogel system for AF repair.⁸⁸ Chemical adsorption to AF tissue is achieved through EDC cross-linking chemistry, which conjugates macromers and proteins via reactive carboxylic acid groups. Four hydrogel formulations were assessed for downstream experimentation with varying macromer to carbodiimide cross-linker stoichiometries: (1) gelatin and γPGA at a 40:1 ratio with EDC, (2) gelatin and γPGA at a 10:1 ratio with EDC, (3) gelatin at a 10:1 ratio with EDC, and (4) γPGA at a 10:1 ratio with EDC. For formulations containing gelatin and γPGA, the carboxylic moieties react with EDC to form a disubstituted o-acylisourea intermediate per repeat unit. The o-acylisourea active ester subsequently reacts with primary amines on the gelatin backbone and primary amines on the collagen backbone in AF matrix to create a covalently linked bridge between the hydrogel and AF. Through EDC chemistry, gelatin polypeptides and γPGA macromers are simultaneously cross-linked to form a stable copolymer hydrogel network. For formulations composed of gelatin without γPGA, EDC reacts with carboxylic groups on the gelatin polypeptide to form an unstable o-acylisourea active ester that forms a stable amide bond with AF collagen. Hydrogel polymerization simultaneously occurs through EDC-based cross-linking between primary amines and carboxylic acid groups across gelatin polypeptides. For formulations composed of γPGA without gelatin, EDC reacts with carboxylic groups on γPGA to form an unstable o-acylisourea active ester that forms a stable amide bond with AF collagen. Hydrogel polymerization simultaneously occurs through EDC-based cross-linking between primary amines and carboxylic acid groups across γPGA macromers (Figure 13).

Figure 13.

Integration mechanism EDC-cross-linked gelatin/γPGA hydrogels.

5.9. Fibrin Hydrogels.

Buser and colleagues investigated the use of the BIOSTAT BIOLOGX fibrin sealant for AF repair and characterized biomechanical and biological effects of repair in a surgically damaged porcine IVD model.⁹⁰ A different fibrin gel formulation was investigated by Zhou et al. as a delivery vehicle for CCL-5 chemoattractant to stimulate endogenous AF repair.⁸⁹ Chemical adsorption of fibrin to AF tissue is achieved through tissue transglutaminase activity (i.e., transglutaminase 2 or “TG2”), where calcium-activated TG2 enzymatically forms covalent cross-links between fibrin’s αC domain and lysine residues on cellular integrin transmembrane receptors or matrix proteins via isopeptide bond formation. Mechanistically, cysteine residues within TG’s active site contain reactive thiol groups that attack carboxamide groups of glutamine residues located on the fibrin backbone. After this first reaction step, the thioester intermediate then reacts with amine groups, which are typically lysine residues found on cellular transmembrane proteins such as integrins or extracellular matrix proteins.^116,117 The stable product following this second reaction step is a cross-linked complex between the fibrin hydrogel, resident cells, and matrix proteins all residing in AF tissue (Figure 14). This cross-linking reaction at the tissue–hydrogel interface simultaneously occurs alongside thrombin-mediated polymerization of fibrinogen monomers once the sealant is injected into the disc space.

Figure 14.

Integration mechanism for fibrin hydrogels.

5.10. PEGDA/Fibrin Hydrogels with CSMA or HAMA Aldehyde.

DiStefano and colleagues developed a two-part AF repair strategy consisting of an oxidized and methacrylated glycosaminoglycan (i.e., dual-modified GAG) and void-filling hydrogel.⁹² The hydrogel is an interpenetrating network composed of a synthetic poly(ethylene glycol) diacrylate (PEGDA) network and natural fibronectin-conjugated fibrin network. Chemical adsorption of the injectable hydrogel is mediated by the dual-modified GAG and forms a covalent bridge between extracellular matrix proteins and the synthetic network of the hydrogel. Aldehyde moieties on oxidized and methacrylated chondroitin sulfate (CSMA aldehyde) or hyaluronic acid (HAMA aldehyde) bond to primary amines of extracellular matrix proteins via Schiff base formation, resulting in two covalent double bonds with ECM proteins per GAG repeat unit. In the presence of ammonium persulfate (APS) and N,N,N′,N′-tetramethylethylenediamine (TEMED) redox initiators, the pendant methacrylate groups on CSMA aldehyde or HAMA aldehyde covalently cross-link with the hydrogel’s PEGDA network once the prepolymer solution is injected into the AF defect space, thus forming the covalent bridge between the IVD and hydrogel. AF repair with this strategy follows a two-step workflow, where AF defect surfaces are first coated with CSMA aldehyde or HAMA aldehyde solution for 5 min. Following 5 min of application, the dual-modified GAG solution is aspirated off of the AF tissue and the prepolymer solution of the interpenetrating network hydrogel is injected into the defect space, where gelation and covalent bonding to IVD matrix proteins occurs simultaneously. (Figure 15) The fibronectin-conjugated fibrin polymer network can also chemically adsorb to IVD tissue through the mechanisms specified in section 5.9.

Figure 15.

Integration mechanism for PEGDA/FN–fibrin hydrogels with CSMA aldehyde or HAMA aldehyde.

6. FUNCTIONAL REGENERATIVE OUTCOMES

Here, we report regenerative outcomes for each hydrogel formulation that has been functionally evaluated using ex vivo or in vivo preclinical models. Although there may be heterogeneity in specific outcome measures across studies, we aggregated regenerative outcomes among four common classes: biomechanical outcomes, biochemical outcomes, radiographic outcomes, and histological outcomes. Following data classification, we then identified hydrogel formulations that were most comprehensively assessed for a given outcome type or identified the most commonly reported output measure(s) within a given category, underscoring their roles as functional metrics of success for AF tissue engineering strategies.

6.1. Biomechanical Outcomes.

Biomechanical analyses included a variety of output measurements across all studies and reported outcomes that (1) determine the effect of repair on axial, torsional, and viscoelastic motion segment responses, (2) assess the durability of a repair strategy, or (3) evaluate mechanical feasibility and corresponding herniation risk after in situ application. Biomechanical performance of AF repair hydrogels is provided in Table 2. Of the 18 hydrogels formulated for AF repair, genipin-cross-linked fibrin and riboflavin-cross-linked collagen hydrogels were the most extensively characterized with regards to biomechanical responses following treatment. When genipin-cross-linked fibrin and riboflavin-cross-linked collagen were compared, output measures were distinctly different from one another; studies analyzing the biomechanical performance of IVDs treated with genipin-cross-linked fibrin quantified axial and torsional responses to physiological loading as well as failure properties, whereas studies analyzing IVDs treated with riboflavin-cross-linked collagen measured hydraulic permeability and stress relaxation responses. Twelve studies assessing biomechanical responses following AF treatment for nine other hydrogel formulations either reported a subset or completely different output measures than those for genipin-cross-linked fibrin and riboflavin-cross-linked collagen, highlighting the large heterogeneity in biomechanical variables reported across studies. This variation in selected output measures makes it increasingly difficult to compare the effectiveness of AF repair across hydrogel compositions. Five of the 18 hydrogels evaluated for AF repair had no reported outcomes with regards to biomechanical responses to treatment, which was a motivation for our summary of functional evaluation methods (Figure 5).

Table 2.

Biomechanical Outcomes for Hydrogels Evaluated in Situ for AF Repair

Hydrogel composition	Abbreviation	Test method of application (Species)	Experimental methods	Outcome	Refs
genipin-cross-linked fibrin	FibGen	ex vivo (bovine and ovine)	cyclic axial/torsional testing	repair was not different than injured controls with respect to torsional range of motion, stiffness, hysteresis, and neutral zone	62–73
			ASTM2255–05 (lap shear)	repair achieves partial restoration of axial properties such as compressive stiffness and neutral zone length in some studies but not all
			ramp-to-failure test	increase in genipin concentration increases motion segment failure strength and decreases herniation risk
		in vivo (ovine)	a	a	73
riboflavin-cross-linked collagen	RF-collagen	ex vivo (rat and ovine)	cyclic axial/torsional testing	repair can be applied to AF defects of varying size (i.e., small and large) and partially restore stress relaxation properties to intact levels, depending on hydrogel’s RF and collagen concentration	74–77
			ASTM2255–05 (lap shear)	repaired IVDs demonstrated hydraulic permeability that was not different than intact IVD controls in a rat-tail model but was higher in an ovine model
			stress relaxation testing	RF–collagen treatment restores torsional stiffness and torque range to intact levels only when used with HA in composite repair strategy
				RF–collagen treatment leads to decrease in tensile stiffness with or without HA composite
		in vivo (rat and ovine)	a	a	77–82
Rose Bengal-cross-linked collagen	RB-collagen	ex vivo (rabbit)	cyclic axial testing	hydrogel constructs were intact and retained in the repair site after 40 320 cycles of compressive and torsional loading	83
			ramp-to-failure test	treatment with hydrogel plug led to higher retention of MSC-collagen microspheres in repair site
		in vivo (rabbit)	a	a
poly(ethylene glycol)–poly (trimethylene carbonate)–hexamethylene diisocyanate	PEG400-TMC3-HDI	ex vivo (bovine and caprine)	cyclic axial/torsional testing	repair led to restoration of axial range of motion, torsional stiffness, torque range, and torsional hysteresis area compared to intact IVD controls	84,85
			ramp-to-failure test	hydrogel-treated IVDs were not different than injured controls with respect to IVD failure strength and subsidence-to-failure
				hydrogel herniated out of repair site within 4 days in bovine organ culture
albumin/glutaraldehyde	BioGlue	ex vivo (bovine)	cyclic axial/torsional testing	repair did not improve axial or torsional motion segment mechanical properties compared to injured controls	69
			ramp-to-failure test	majority of repaired motion segments failed during torsional loading protocol
n-butyl cyanoacrylate (n-BCA) and octyl cyanoacrylate (OCA)	cyanoacrylate/LiquiBand	ex vivo (porcine)	ramp-to-failure test	repair with cyanoacrylate glue led to a significant increase in axial failure load compared to untreated IVDs	86
		in vivo (porcine)	a	a
citric acid–1-ethyl-3-(3-dimethylaminopropyl)carbodiimide/N-hydroxysuccinimide type I collagen	CA-EDC/NHS collagen	in vivo (rat)	a	a	87
1-(3-dimethylaminopropyl)-3-ethyl-carbodiimide hydrochloride-cross-linked gelatin/poly(γ-glutamic acid)	EDC-gelatin-γPGA	ex vivo (bovine)	quantitative discomanometry	testing indicated that AF treatment with a hydrogel formulation of 10% gelatin, 3% γPGA, and 2% EDC achieved leakage and saturate pressures that were not different than intact IVDs and were significantly greater than injured controls	88
fibrin	fibrin	ex vivo (bovine and porcine)	quantitative discomanometry	repair led to a significant increase in relative pressure modulus and relative failure pressure compared to injured controls at 6- and 12-weeks postop in an in vivo porcine model	89,90
		in vivo (porcine and ovine)	a	a
poly(lactic-co-glycolic acid)/fibrin	PLGA/fibrin	in vivo (rabbit)	a	a	91
poly(ethylene glycol) diacrylate + chondroitin sulfate or hyaluronic acid methacrylate aldehyde	PEGDA/fibrin + CSMA or HAMA aldehyde	ex vivo (bovine)	ASTM2255–05(lap shear)	treatment of AF tissue with CSMA aldehyde and HAMA aldehyde led to significant increase in hydrogel adhesion strength	92
			ramp-to-failure test	higher molecular weight PEGDA resulted in a lower IVD herniation risk
				matched herniation risk compared to the current standard of care
poly(lactic-co-glycolic acid)	PLGA	in vivo (rabbit)	a	a	93
alginate/collagen	alginate/collagen	ex vivo (bovine)	stress relaxation testing	incorporation of collagen in alginate did not significantly affect peak stress, aggregate compressive equilibrium modulus, and swelling ratio of hydrogels	94
alginate	alginate	ex vivo (bovine)	stress relaxation testing	hydrogels cultured in hypoxic and normoxic conditions demonstrated stress relaxation behavior that was not significantly different than native AF tissue	95
				TGF-β3 incorporation in normoxic culture led to peak stresses that were not significantly different than native AF values, although all equilibrium moduli were significantly lower than AF values regardless of oxygen tension or TGF-β3 incorporation
hyaluronic acid	HA	ex vivo (bovine)	a	a	96
		in vivo (rat)	a	a	97
hyaluronic acid/albumin	HA/albumin	in vivo (ovine)	a	a	98
ultrapurified alginate	UPAL	ex vivo (ovine)	cyclic axial/torsional testing	no hydrogel protrusion was observed after static and dynamic mechanical loading	99
				repair led to partial restoration of IVD compressive stiffness and was not significantly different than intact IVD controls
		in vivo (rabbit and ovine)	a	a
cellulose nanofiber-reinforced chitosan	CNF/CHI	ex vivo (porcine)	ramp-to-failure test	repaired IVDs demonstrated greater motion segment ultimate strength compared to injured controls and was above physiological stress levels	100
				slope of linear elastic region was greater in repaired IVDs than injured controls
				ultimate strength and slope of the linear elastic region were less than that of intact IVDs

^aN/A = Not applicable.

Notably, eight formulations led to improvements in IVD failure properties or retention of the hydrogel after biomechanical loading, indicating functional success of hydrogel integration and promise for clinical utility. To relate these outcomes to their mechanisms of integration, six formulations (FibGen, RF-collagen, RB-collagen, cyanoacrylate, EDC-gelatin-γPGA, and fibrin) primarily interact with the AF by means of chemisorption and two formulations (UPAL and CNF/CHI) primarily interact with the AF by means of electrostatic interactions, physisorption, and interdiffusion. These outcomes suggest that specific mechanisms of covalent bonding achieve the greatest likelihood of biomechanical success by mitigating the risk of construct herniation. Covalent bonding mediated by natural compounds (genipin or transglutaminase), UV-catalyzed photoinitiation (riboflavin or rose bengal), or spontaneous Michael addition demonstrated the greatest promise for AF repair hydrogels. Although covalent bonding generally led to successful outcomes with respect to biomaterial retention and herniation risk, it is important to note that not all mechanisms of covalent bonding achieve this type of functional success. Specifically, two hydrogel formulations that employ covalent bonding mechanisms to integrate with the AF (albumin–glutaraldehyde and PEG₄₀₀-TMC3-HDI) led to hydrogel protrusion from the repair site at some point during biomechanical testing or organ culture. For albumin–glutaraldehyde hydrogels, covalent bonding with AF tissue occurs via Schiff base formation between aldehyde moieties on the terminal ends of glutaraldehyde and primary amines on ECM proteins and albumin. For PEG₄₀₀-TMC3-HDI hydrogels, covalent bonding with AF tissue occurs via carbamide formation between HDI groups on the terminal end of the macromer and primary amines on ECM proteins. To improve the biomechanical performance of these two formulations, there could be additional biochemical optimizations that tune biomaterial bonding strength so as to achieve functional success and mitigate the risk of herniation. Taken together, these findings underscore the critically important role that integration mechanisms serve for AF repair hydrogels and highlight the need for tissue engineers to understand these mechanisms as they advance candidate formulations through the preclinical pipeline.

6.2. Biochemical and Biological Outcomes.

Biochemical and biological assessments provide important quantitative insight with respect to maintenance or changes in cellular phenotype, anabolism, catabolism, cytokine production, and cell proliferation or apoptosis. Biochemical assays that quantitatively measure GAG content and collagen content are useful to inferentially determine whether these hydrogels promote anabolic or catabolic behavior of resident cells within the IVD or exogenously delivered cells. To ensure the longevity and functional success of repair, these biomaterials should not elicit significant catabolic activity in order to preserve the integration strength of the implantable constructs. Of the 18 hydrogel formulations, studies for seven candidate materials employ biochemical assays to evaluate ECM content. Fibrin was the only formulation that demonstrated a significantly higher amount of GAG compared to injured controls and also matched that of the intact IVD. Studies corresponding to the other six formulations (genipin-cross-linked fibrin, PLGA/fibrin, alginate, hyaluronic acid, and hyaluronic acid/albumin) demonstrate either no change in ECM content over time or a significantly lower amount of GAG or collagen content compared to injured or intact controls. In addition to analytical measurements of ECM content, biological assessments such as cellular viability, proliferation, and gene expression analyses elucidate the effect of biomaterial implantation on cellular health and phenotype. Studies for five of 18 hydrogels (cyanoacrylate, alginate, alginate/collagen, and ultrapurified alginate) assessed cellular viability, where alginate-based hydrogels supported high levels of cell viability and cyanoacrylate-based adhesives led to a decrease in cell viability. Studies pertaining to another four formulations (PEG₄₀₀-TMC3-HDI, CA-EDC/NHS collagen, alginate, and alginate/collagen) examined proliferative activity, where each of these four gels were determined to be permissive of cell proliferation. With respect to phenotype, studies corresponding to five of 18 hydrogels (FibGen, PLGA, PLGA/fibrin, HA, and HA/albumin) reported fold changes in gene expression via qPCR in response to repair, where gene sets significantly varied across all studies. Biochemical and biological output measures and their corresponding outcomes for all AF repair hydrogels are provided in Table 3.

Table 3.

Biochemical and Biological Outcomes for Hydrogels Evaluated in Situ for AF Repair

Hydrogel composition	Abbreviation	Test method of application (Species)	Experimental methods	Outcome	Refs
genipin-cross-linked fibrin	FibGen	ex vivo (bovine and ovine)	calcein AM/DAPI viability assay	analytical validation of chemical cross-linking and structure	62–73
			1H NMR	incorporation of genipin or cell adhesion molecules to fibrin gel did not affect COLI, MMP13, and IL-1β expression
			ninhydrin assay	no difference was observed between injured and repaired IVDs in organ culture with respect to ACAN, COL1, COL2, BGN, MMP3, MMP12, ADAMTS4, IL-1β, IL-8, CCL2, C0X2, and NGF expression
			qRT-PCR	no significant changes in GAG content for cell-laden hydrogels at 7-, 28-, and 49-days in culture
		in vivo (ovine)	b	b	73
riboflavin-cross-linked collagen	RF-collagen	ex vivo (rat and ovine)	b	b	74–77
		in vivo (rat and ovine)	b	b	77–82
Rose Bengal-cross-linked collagen	RB-collagen	ex vivo (rabbit)	FTIR	RB-concentration dependent changes in amide I, II, and III peaks, suggesting cross-linking and tissue bonding is improved with an increase in RB concentration	83
		in vivo (rabbit)	b	b
poly (ethylene glycol)–poly (trimethylenecarbonate)–hexamethylene diisocyanate	PEG400-TMC3-HDI	ex vivo (bovine and caprine)	QuantiFluor dsDNA assay	hydrogel was more permissive of AF cell proliferation than Dermabond as demonstrated by an increase in DNA concentration over a 7-day period	84,85
albumin/glutaraldehyde	BioGlue	ex vivo (bovine)	b	b	69
n-butyl cyanoacrylate (n-BCA) and octyl cyanoacrylate (OCA)	cyanoacrylate/LiquiBand	ex vivo (porcine)	XTT assay	MSC viability significantly decreased at 1 and 4 days of exposure to the bio material adhesive	86
		in vivo (porcine)	b	b
citric acid-1-ethyl-3-(3-(dimethylamino)propyl) carbodiimide/N-hydroxysuccinimide type I collagen	CA-EDC/NHScollagen	in vivo (rat)	MTT assay	no difference in L929 cellular proliferation between all hydrogel formulations and control IVD	87
1-(3-(dimethylamino)propyl)-3-ethyl-carbodiimide hydrochloride-cross-linked gelatin/poly(γ-glutamic acid)	EDC-gelatin-γPGA	ex vivo (bovine)	b	b	88
fibrin	fibrin	ex vivo (bovine and porcine)	qRT-PCR	no change in ADAMTS4, TIMP1, MMP1, MMP12, TNFα, or COL1 gene expression after treatment with CCL5	89,90
		in vivo (porcine and ovine)	DMMB assay	relative GAG content was significantly higher than injured controls and within range of control group levels at 12 weeks postop
			Pico Green assay ELISA	ELISA indicated attenuation of TNFα and IL-1β levels and increase of TGF-β levels in repaired IVDs compared to injured IVDs 2–3 weeks postop
poly(lactic-co-glycolic acid)/fibrin	PLGA/fibrin	in vivo (rabbit)	DMMB assay	no difference in ACAN, COL2A2, COL1A1, and MMP3 expression between repair group and intact control group	91
			qRT-PCR	repair showed significant increase in ACAN and COL2A2 and a significant decrease in COL1A1 and MMP3 expression levels compared to injury groups
				repair demonstrated significant decrease in sGAG content 6 months postop compared to all groups
poly(ethylene glycol) diacrylate + chondroitin sulfate or hyaluronic acid methacrylate aldehyde	PEGDA/fibrin + CSMA or HAMA aldehyde	ex vivo (bovine)	1H NMR	methacrylation and oxidation analytically quantified for all CSMA aldehyde and HAMA aldehyde formulations screened	92
			CellTiter-Glo 2.0 cell viability assay	no difference in AF cell viability between live cell control and HAMA aldehyde-treated cells
			TNBS assay
poly(lactic-co-glycolic acid)	PLGA	in vivo (rabbit)	qRT-PCR	no difference in ACAN, COLI, COL2, MMP3 expression 1, 3, and 6 months postop between repaired IVDs and intact controls	93
alginate/collagen	alginate/collagen	ex vivo (bovine)	hydroxyproline assay	concentration-dependent increase in collagen content upon incorporation in the hydrogel	94
			Hoechst Bisbenzimide 33258 dye assay	incorporation of collagen and TGF-β3 led to an increase in DNA content at 2-, 10-, and 21-days in culture compared to unloaded and alginate-only controls
			LIVE/DEAD assay	high MSC viability in hydrogels up to 5 weeks
alginate	alginate	ex vivo (bovine)	DMMB assay	sGAG content (normalized to DNA abundance) was not different than AF for hydrogels that incorporated TGF-β3 and cultured in 5% O2	95
			hydroxyproline assay	incorporation of TGF-β3 led to higher DNA content in hydrogels after 21 days in hypoxic and normoxic culture compared to day 0
			Hoechst Bisbenzimide 33258 dye assay	collagen content (normalized to DNA abundance) was significantly lower than AF for all constructs, regardless of oxygen tension or TGF-β3 incorporation
			LIVE/DEAD assay	high AF cell viability after 21 days in culture across hypoxic and normoxic conditions irrespective of TGF-β3 incorporation
hyaluronic acid	HA	ex vivo (bovine)	qRT-PCR	repair led to downregulation of IFNAR1, IFNAR2, STAT1, STAT2, IFIT3, ADAMTS4, and IGFBP3 under IFNα challenge	96
				repair led to an upregulation of ACAN and C0L1A2 with and without IFNα challenge
		in vivo (rat)	high-performance liquid chromatography	repair led to temporally significant differences in C0S and C4S disaccharide content in the AF over a 56-day period	97
hyaluronic acid/albumin	HA/albumin	in vivo (ovine)	DMMB assay	cell-laden and acellular hydrogel repairs did not have an effect on IVD GAG, DNA, and collagen content compared to injured and intact IVD controls	98
			hydroxyproline assay	acellular hydrogels exhibited higher levels of C0L1A2 and lower levels of IL-1β expression compared to cell-laden hydrogels
			Pico Green assay qRT-PCR	repair resulted in elevated C0L2A1, ACAN, and PRG4 expression, and lower C0L1A2, COL10, and IL-1β expression
ultrapurified alginate	UPAL	ex vivo (human and ovine)	flow cytometry	hydrogels can support 3D cell culture with high viability and low apoptosis (~75% PI and Annexin V-negative cells) for up to 48 h	99
		in vivo (rabbit and ovine)	b	b
cellulose nanofiber-reinforced chitosan	CNF/CHI	ex vivo (porcine)	b	b	100

^aExperimental method abbreviations: ¹H NMR, proton nuclear magnetic resonance spectroscopy; qRT-PCR, quantitative real-time polymerase chain reaction; FTIR, Fourier transform infrared spectroscopy; DMMB, dimethylmethylene blue; MTT, 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide; XTT, 2,3-bis(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium-5-carboxanilide; TNBS, 2,4,6-trinitrobenzenesulfonic acid. Gene abbreviations: ACAN, aggrecan; COL1A1, collagen type I alpha 1; COL2A1, collagen type II alpha 1; COL10, collagen type X; MMP3, matrix metalloproteinase-3; MMP13, matrix metalloproteinase-13; IL-1β, interleukin-1 beta; BGN, biglycan; ADAMTS4, A disintegrin and metalloproteinase with thrombospondin motifs 4; IL-8, interleukin-8; CCL2, C–C motif chemokine ligand 2; COX2, cyclooxygenase-2; NGF, nerve growth factor; IFNAR1, interferon alpha and beta receptor subunit 1; IFNAR2, interferon alpha and beta receptor subunit 2; STAT1, signal transducer and activator of transcription 1; STAT2, signal transducer and activator of transcription 2; IFIT3, interferon induced protein with tetratricopeptide repeats 3; IGFBP3, insulin-like growth factor-binding protein 3.

^bNot applicable.

6.3. Radiographic Outcomes.

Radiography allows investigators to examine spinal anatomy after repair and quantify parameters that are strongly associated with painful degeneration to determine if AF repair can offset the degenerative progression that is known to occur after discectomy.^10,118 Magnetic resonance imaging (MRI) provides insight with respect to NP hydration state, end plate changes according to Modic classification criteria, and disc degeneration grade using unmodified and modified Pfirrmann grading systems.^119–122 X-ray imaging enables researchers to detect osteophyte formation and carry out morphometric analyses in order to quantify important parameters that are strongly correlated with degeneration-associated pain, including disc height (DH) and/or disc height index (DHI).¹²³ Of the 18 formulations, studies pertaining to four hydrogels (riboflavin-cross-linked collagen, CA-EDC/NHS collagen, PLGA, and PLGA/fibrin) comprehensively assess both DH/DHI and Pfirrmann grade, whereas studies for six other formulations quantify DH/DHI via either X-ray or MRI-based parameters. Riboflavin-cross-linked collagen (cell-laden), CA-EDC/NHS collagen, PLGA, and PLGA/fibrin formulations demonstrated significant improvements in DHI and degeneration grade compared to injured IVDs, suggesting that partial restoration of functionally important parameters can be achieved with these hydrogel formulations. Notably, studies for seven candidate hydrogels did not incorporate any radiographic analyses to determine changes in DH/DHI or MRI-based output measures in response to hydrogel repair. Six of these seven hydrogels did not have any associated radiographic outcomes because these repair strategies were strictly evaluated in an ex vivo IVD model. LiquiBand was the only adhesive of these seven formulations that included a study with an in vivo application of the repair strategy without any reported radiographic outcomes. In this in vivo study, Kang et al. exclusively performed histological assessments of LiquiBand in a porcine repair model following discectomy. Radiographic outcomes for all AF repair hydrogels are provided in Table 4.

Table 4.

Radiographical Outcomes for Hydrogels Evaluated in Situ for AF Repair

Hydrogel composition	Abbreviation	Test method of application (Species)	Experimental methods	Outcomes	Refs
genipin-cross-linked fibrin	FibGen	ex vivo (bovine and ovine)	X-ray	repaired IVDs did not demonstrate any difference in disc height loss compared to intact controls	62–73
		in vivo (ovine)	T2-weighted MRI	Pfirrmann grade was not different between injured and treated IVDs and were both lower than intact controls	73
				DHI was not different between injured and treated IVDs and were both lower than intact controls
riboflavin-cross-linked collagen	RF-collagen	ex vivo (rat and ovine)	T2-weighted MRI	IVDs treated with RF-collagen demonstrated significantly lower NP volume and T2 relaxation time compared to intact controls	74–77
				IVD treatment with RF-collagen and HA composite repair strategy restored T2 relaxation time to intact controls
		in vivo (rat and ovine)	T2-weighted MRI	higher RF concentration (0.5 mM) led to a higher DHI and NP voxel count than lower RF concentrations (0 and 0.25 mM) over 5 weeks postop in a rat-tail model	77–82
			X-ray	RF-collagen with and without HA composite demonstrated no difference in Pfirrmann grade or DHI compared to intact IVDs at 6-weeks postop in a sheep model
				cell-laden repair groups demonstrated significantly higher DHI (relative to intact) and T2 relaxation time (normalized to intact) at 6-weeks postop in a sheep model
				repaired IVDs with cell-laden hydrogels had a lower degenerative grade than acellular gel and injured controls according to the Pfirrmann scoring system
Rose Bengal-cross-linked collagen	RB-collagen	ex vivo (rabbit)	a	a	83
		in vivo (rabbit)	T2-weighted MRI	repair led to a significant reduction in the formation of osteophytes
			micro CT
			X-ray
poly(ethylene glycol)–poly(trimethylenecarbonate)–hexamethylene diisocyanate	PEG400-TMC3-HDI	ex vivo (bovine and caprine)	a	a	84,85
albumin/glutaraldehyde	BioGlue	ex vivo (bovine)	X-ray	repaired IVDs demonstrated a significant increase in disc height compared to injured IVDs	69
n-butyl cyanoacrylate (n-BCA) and octyl cyanoacrylate (OCA)	cyanoacrylate/LiquiBand	ex vivo (porcine)	a	a	86
		in vivo (porcine)	a	a
citric acid-1-ethyl-3-(3-(dimethylamino)propyl)carbodiimide/N-hydroxysuccinimide type I collagen	CA-EDC/NHS collagen	in vivo (rat)	X-ray	higher CA concentrations led to an increase in NP voxel count and less overall decrease in DHI compared to lower CA concentrations and injured controls over a 4-week period in an in vivo rat model	87
			T2-weighted MRI
1-(3-(dimethylamino)propyl)-3-ethyl-carbodiimide hydrochloride-cross-linked gelatin/poly(γ-glutamic acid)	EDC-gelatin-γPGA	ex vivo (bovine)	a	a	88
fibrin	fibrin	ex vivo (bovine and porcine)	a	a	89,90
		in vivo (porcine and ovine)	T2-weighted MRI	no difference in Pfirrmann degeneration grade was observed between repaired and injured IVDs
poly(lactic-co-glycolic acid)/fibrin	PLGA/fibrin	in vivo (rat)	T2-weighted MRI	DHI was significantly lower in repair and injury groups compared to intact controls up to 6 months postop	91
				DHI of repair group was significantly higher than all injury groups through 6 months postop
				MRI grade was significantly higher than intact controls and significantly lower than injury groups at 6 months postop
poly(ethylene glycol) diacrylate + chondroitin sulfate or hyaluronic acid methacrylate aldehyde	PEGDA/fibrin + CSMA or HAMA aldehyde	ex vivo (bovine)	a	a	92
poly(lactic-co-glycolic acid)	PLGA	in vivo (rabbit)	X-ray	repaired IVDs demonstrated less of a decrease in DHI than injury groups at 0.5-, 1-, 3-, and 6-months postop	93
			T2-weighted MRI	repaired IVDs had a lower degenerative grade according to the Pfirrmann scoring system at 3- and 6- months postop
alginate/collagen	alginate/collagen	ex vivo (bovine)	a	a	94
alginate	alginate	ex vivo (bovine)	a	a	95
hyaluronic acid	HA	ex vivo (bovine)	a	a	96
		in vivo (rat)	X-ray	repair resulted in a significantly higher DHI compared to sham and injury groups at 7-, 28-, and 56-days postop	97
hyaluronic acid/albumin	HA/albumin	in vivo (ovine)	Tl- and T2- weighted MRI	reduction in NP signal intensity was observed 2- and 4-weeks postop across all repair groups no observable difference between all groups in MRI imaging at 3 months postop	98
ultrapurified alginate	UPAL	ex vivo (human and ovine)	a	a	99
		in vivo (rabbit and ovine)	T2-weighted MRI	repaired IVDs demonstrated a significantly lower Pfirrmann degeneration grade than injury groups 4-, 12-, and 24-weeks postop
				MRI index (NP area × average signal intensity) was significantly higher in repaired IVDs than injured IVDs across all time points
cellulose nanofiber-reinforced chitosan	CNF/CHI	ex vivo (porcine)	a	a	100

^aNot applicable.

6.4. Histological Outcomes.

Histological examination of the IVD following AF repair can provide visual data regarding in situ retention of biomaterials, in vivo degradation, host tissue biocompatibility, and integration of hydrogels applied to AF defects. These assessments generally rely on tinctorial staining to visualize the architecture, quality, and morphological integrity of AF tissue and biomaterials after repair. Immunohistochemical analyses also enable investigators to probe specific changes in cellular protein expression, where immunohistochemical markers varied across studies to examine biological repair responses to a given hydrogel formulation. Studies for seven hydrogel formulations histologically examined biomaterial integration, retention, adherence, and resorption, where all seven formulations demonstrated some degree of biomaterial degradation and partial adherence to AF tissue after in vivo implantation. Additionally, tinctorial staining of repaired IVDs generally showed that hydrogel treatment leads to better preserved IVD architecture with centrally retained NP tissue, indicating some level of success for these hydrogels to serve as an injectable adhesive and tissue sealant. Immunohistochemical (IHC) analyses were reported for five hydrogel formulations (genipin-cross-linked fibrin, PLGA/fibrin, HA, HA/albumin, and ultrapurified alginate), where COL1 was the most commonly used marker for IHC assessments. All reports that investigated COL1 expression demonstrated maintenance of AF phenotype after repair or cellular encapsulation by COL1 immunopositivity within or adjacent to the hydrogel repair. Histological outcomes for all AF repair hydrogels are provided in Table 5.

Table 5.

Histological Outcomes for Hydrogels Evaluated in Situ for AF Repair

Hydrogel composition	Abbreviation	Test method of application (Species)	Experimental methods	Outcome	Refs
genipin-cross-linked fibrin	FibGen	ex vivo (bovine and ovine)	IHC	Ki67 expression did not change with a change in fibrin or genipin concentration	62–73
			tinctorial staining	COL1 expression increased with a decrease in both fibrin and genipin concentration
				hydrogel was retained within defect site after mechanical loading
		in vivo (ovine)	tinctorial staining	partial degradation and cell infiltration occurs by 8 weeks and complete resorption occurs by 16 weeks	73
riboflavin-cross-linked collagen	RF-collagen	ex vivo (rat and ovine)	tinctorial staining	hydrogel treatment decreased NP height loss and promoted reorganization of lamellar layers	74–77
			SHG	repair with collagen AF patch decreased buckling and inversion of AF lamellae
				collagen gel demonstrated significant binding with AF tissue after ChABC treatment
		in vivo (rat and ovine)	tinctorial staining	partial adherence of hydrogel to AF tissue and some biomaterial degradation observed at 2- and 5-weeks in a rat-tail model	77–82
			SHG	repair leads to centrally retained NP tissue and formation of fibrous capsule at outer AF after 5 weeks in a rat-tail model
				acellular and MSC-laden repairs led to hypointense and hyperintense staining, respectively, at the outer AF 6-weeks postop in a sheep model
				hydrogel treatment led to increased aligned collagen in the direction of lamellar bundles in a sheep model
Rose Bengal-cross-linked collagen	RB-collagen	ex vivo (rabbit)	IHC	injection portal was evidence 1-month postop	83
			tinctorial staining	no osteophyte formation in repaired IVDs
		in vivo (rabbit)	a	a
poly(ethylene glycol)–poly(trimethylenecarbonate)–hexamethylene diisocyanate	PEG400-TMC3-HDI	ex vivo (bovine and caprine)	tinctorial staining	mechanical failure predominantly occurred at the interface between hydrogel and AF	84,85
albumin/glutaraldehyde	BioGlue	ex vivo (bovine)	a	a	69
n-butyl cyanoacrylate (n-BCA) and octyl cyanoacrylate (OCA)	cyanoacrylate/LiquiBand	ex vivo (porcine)	a	a	86
		in vivo (porcine)	IHC	H&E and Picrosirius Red staining indicated signs of inflammatory reaction and fibrous scar tissue
			tinctorial staining
				αSMA IHC indicated blood vessel ingrowth after repair
citric acid-l-ethyl-3-(3-dimethylaminopropyl)carbodiimide/N-hydroxysuccinimide type I collagen	CA-EDC/NHScollagen	in vivo (rat)	tinctorial staining	repaired IVDs demonstrated a CA-dependent decrease in histological grading score compared to untreated IVD controls	87
				H&E and Saf-O staining indicated a gel-like material was able to bridge the AF together and close the needle track
1-(3dimethylaminopropyl)-3-ethyl-carbodiimide hydrochloride-cross-linkedgelatin/poly(γ-glutamic acid)	EDC-gelatin-γPGA	ex vivo (bovine)	a	a	88
fibrin	fibrin	ex vivo (bovine and porcine)	tinctorial staining	histopathological scoring was not different between injured and repaired IVDs with and without CCL5	89,90
		in vivo (porcine and ovine)		repaired IVDs show improvement in AF architecture compared to injured controls and retention of NP tissue up to 12 weeks postop
				preservation of NP volume over 12 weeks
poly(lactic-co-glycolic acid)/fibrin	PLGA/fibrin	in vivo (rat)	IHC	significant biomaterial resorption at 6 months postop with clusters of new repair tissue around remaining hydrogel	91
			tinctorial staining	PGP 9.5 IHC indicated that nerve tissue extended into the inner AF region, with proximity to blood vessels and found in the generated repair tissue in the hydrogel treatment group
				Saf-O staining indicated significant decrease in proteoglycan content within the repair site
poly(ethylene glycol) diacrylate + chondroitin sulfate or hyaluronic acid methacrylate aldehyde	PEGDA/fibrin + CSMA or HAMA aldehyde	ex vivo (bovine)	tinctorial staining	Picrosirius Red/Alcian Blue staining indicated that AF tissue and injectable hydrogel are adsorbed via GAG-mediated covalent bonds with contiguous boundaries	92
poly(lactic-co-glycolic acid)	PLGA	in vivo (rabbit)	tinctorial staining	immature repair tissue observed at 6 months postop	93
				significant biomaterial degradation after 6 months
alginate/collagen	alginate/collagen	ex vivo (bovine)	tinctorial staining	TGF-β3 incorporation enhanced collagen elaboration	94
alginate	alginate	ex vivo (bovine)	IHC	significant collagen and proteoglycan content deposited after treatment with AF cell-laden hydrogels	95
			tinctorial staining
hyaluronic acid	HA	ex vivo (bovine)	IHC	hydrogel treatment decreases IFIT3, IGFBP3 and Casp3 expression	96
				hydrogel treatment increases ACAN and COL1 expression
		in vivo (rat)	IHC	hydrogel treatment decreases IGFBP3 and Casp3 expression	97
				hydrogel treatment decreases WGA, SNA-I, MAA, CON-A binding
				hydrogel treatment increases PNA binding
				hydrogel treatment increases ACAN and HAPLN1 expression
hyaluronic acid/albumin	HA/albumin	in vivo (ovine)	IHC	NP and AF similarity observed between repaired and intact PVDs	98
			tinctorial staining	ACAN, COL1, and COL2 IHC showed maintenance of phenotype after repair
ultrapurified alginate	UPAL	ex vivo (human and ovine)	a	a	99
		in vivo (rabbit and ovine)	IHC	repaired IVDs underwent less degeneration than discectomy group 4-, 12-, and 24-weeks postop
			tinctorial staining	COL2 IHC was not different between repaired and intact ovine IVDs
cellulose nanofiber-reinforced chitosan	CNF/CHI	ex vivo (porcine)	a	a	100

^aNot applicable.

7. CONCLUSIONS

In this systematic review, we provide a comprehensive list of polymeric biomaterials that have been utilized in hydrogel systems for AF repair and propose their mechanisms of tissue integration as well as their regenerative outcomes. By elucidating the adhesive mechanisms for previously developed hydrogels, we hope to catalyze the development of next generation AF repair strategies that investigate new biochemical modifications of natural and synthetic polymers and exploit any number of the proposed surface chemistries to enhance biomolecular interactions at the AF–hydrogel interface. When reviewing surface chemistries and functional outcomes of the previously developed hydrogels, we observed large heterogeneity across their interfacing mechanisms and reported output measurements regardless whether the outcomes pertained to biomechanical, biochemical, radiographic, or histological assessments. With respect to clinical utility, adhesion strength and construct herniation risk are the most important design parameters, where covalent bonding strategies were found to be the most successful in mitigating herniation risk and construct protrusion after biomechanical loading. Specifically, formulations that achieve covalent bonding with AF tissue via natural compounds (genipin or transglutaminase), UV-catalyzed photoinitiators (riboflavin or rose bengal), or spontaneous Michael addition were the most promising candidates. While several hydrogels improved IVD biomechanical properties compared to injured IVD controls that emulate the surgical standard of care, none restored properties to the intact condition, motivating the need for new bonding methods or biochemical optimizations for these AF repair strategies. Furthermore, the disparity in outcome measures not only made it difficult to compare the effectiveness of repair across different hydrogel compositions but also motivated us to provide functional evaluation methods to standardize material assessments using in vitro systems and accelerate their translation to preclinical animal models for in situ repair. As regenerative medicine emerges as a prominent option for the treatment of AF defects, we highlight the critical role of interface tissue engineering in the development of these treatment options and aim to facilitate the translation of new hydrogel technologies into the clinic for AF repair.