# Design and clinical application of injectable hydrogels for musculoskeletal therapy

Øystein Øvrebø, Giuseppe Perale, Jonathan Wojciechowski, Cécile Echalier, Jonathan Jeffers, Molly Stevens, Håvard Haugen, Filippo Rossi

► To cite this version:

Øystein Øvrebø, Giuseppe Perale, Jonathan Wojciechowski, Cécile Echalier, Jonathan Jeffers, et al. Design and clinical application of injectable hydrogels for musculoskeletal therapy. Bioengineering & Translational Medicine, 2022, 7 (2), pp.e10295. 10.1002/btm2.10295. hal-04601760

HAL Id: hal-04601760

https://hal.science/hal-04601760

Submitted on 5 Jun 2024

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# Design and clinical application of injectable hydrogels for musculoskeletal therapy

Øystein Øvrebø, Giuseppe Perale, Jonathan Wojciechowski, Cécile Echalier, Jonathan Jeffers, Molly Stevens, Håvard Haugen, Filippo Rossi

► To cite this version:

Øystein Øvrebø, Giuseppe Perale, Jonathan Wojciechowski, Cécile Echalier, Jonathan Jeffers, et al. Design and clinical application of injectable hydrogels for musculoskeletal therapy. Bioengineering & Translational Medicine, 2022, 7 (2), pp.e10295. 10.1002/btm2.10295. hal-04601760

HAL Id: hal-04601760

https://hal.science/hal-04601760v1

Submitted on 5 Jun 2024

HAL is a multi-disciplinary open access
archive for the deposit and dissemination of sci-
entific research documents, whether they are pub-
lished or not. The documents may come from
teaching and research institutions in France or
abroad, or from public or private research centers.

L'archive ouverte pluridisciplinaire HAL, est
destinée au dépôt et à la diffusion de documents
scientifiques de niveau recherche, publiés ou non,
émanant des établissements d'enseignement et de
recherche français ou étrangers, des laboratoires
publics ou privés.

<figure><img src="image_1.png" /></figure>

Distributed under a Creative Commons Attribution 4.0 International License
REVIEW ARTICLE

# Design and clinical application of injectable hydrogels for musculoskeletal therapy

Øystein Øvrebø<fnref n="1" />,<fnref n="2" />,<fnref n="3" /> ID

Cécile Echalier<fnref n="7" />,<fnref n="8" />,<fnref n="9" />,<fnref n="10" /> ID

Håvard J. Haugen<fnref n="2" />,<fnref n="3" /> ID

Giuseppe Perale<fnref n="4" />,<fnref n="5" />,<fnref n="6" /> ID

Jonathan R. T. Jeffers<fnref n="11" /> ID

Filippo Rossi<fnref n="1" /> ID

Jonathan P. Wojciechowski<fnref n="7" />,<fnref n="8" />,<fnref n="9" /> ID

Molly M. Stevens<fnref n="7" />,<fnref n="8" />,<fnref n="9" /> ID

Correspondence

Håvard J. Haugen, Department of Biomaterials, Institute of Clinical Dentistry, University of Oslo, Geitmysrveien 71, 0455 Oslo, Norway.

Email: h.j.haugen@odont.uio.no

Funding information

Aker Scholarship; H2020 Marie Skłodowska-Curie Actions, Grant/Award Number: 801133; Norges Forskningsråd, Grant/Award Number: 327512; UK Research and Innovation, Grant/Award Number: MR/R015651/1

Abstract

Musculoskeletal defects are an enormous healthcare burden and source of pain and disability for individuals. With an aging population, the proportion of individuals living with these medical indications will increase. Simultaneously, there is pressure on healthcare providers to source efficient solutions, which are cheaper and less invasive than conventional technology. This has led to an increased research focus on hydrogels as highly biocompatible biomaterials that can be delivered through minimally invasive procedures. This review will discuss how hydrogels can be designed for clinical translation, particularly in the context of the new European Medical Device Regulation (MDR). We will then do a deep dive into the clinically used hydrogel solutions that have been commercially approved or have undergone clinical trials in Europe or the United States. We will discuss the therapeutic mechanism and limitations of these products. Due to the vast application areas of hydrogels, this work focuses only on treatments of cartilage, bone, and the nucleus pulposus. Lastly, the main steps toward clinical translation of hydrogels as medical devices are outlined. We suggest a framework for how academics can assist small and medium MedTech enterprises conducting the initial clinical investigation and post-market clinical follow-up required in the MDR. It is evident that the successful translation of hydrogels is governed by acquiring high-quality pre-clinical and clinical data confirming the device mechanism of action and safety.

KEYWORDS

bone regeneration, cartilage regeneration, clinical translation, hydrogels, medical devices, regenerative medicine, viscosupplementation

<figure><img src="image_2.png" /></figure>

<footnote n="1">¹Department of Chemistry, Materials and Chemical Engineering “Giulio Natta”, Politecnico di Milano, Milano, Italy</footnote>

<footnote n="2">²Department of Biomaterials, Institute of Clinical Dentistry, University of Oslo, Oslo, Norway</footnote>

<footnote n="3">³Material Biomimetic AS, Oslo Science Park, Oslo, Norway</footnote>

<footnote n="4">⁴Industrie Biomediche Insubri SA, Mezzovico-Vira, Switzerland</footnote>

<footnote n="5">⁵Faculty of Biomedical Sciences, University of Southern Switzerland, Lugano, Switzerland</footnote>

<footnote n="6">⁶Ludwig Boltzmann Institute for Experimental and Clinical Traumatology, Vienna, Austria</footnote>

<footnote n="7">⁷Department of Materials, Imperial College London, London, UK</footnote>

<footnote n="8">⁸Department of Bioengineering, Imperial College London, London, UK</footnote>

<footnote n="9">⁹Institute of Biomedical Engineering, Imperial College London, London, UK</footnote>

<footnote n="10">¹⁰Hybrid Technology Hub, Centre of Excellence, Institute of Basic Medical Science, University of Oslo, Oslo, Norway</footnote>

<footnote n="11">¹¹Department of Mechanical Engineering, Imperial College London, London, UK</footnote>
<figure><img src="image_2.png" /></figure>

REVIEW ARTICLE

# Design and clinical application of injectable hydrogels for musculoskeletal therapy

Øystein Øvrebø<fnref n="1" />,<fnref n="2" />,<fnref n="3" /> ID

Cécile Echalier<fnref n="7" />,<fnref n="8" />,<fnref n="9" />,<fnref n="10" /> ID

Håvard J. Haugen<fnref n="2" />,<fnref n="3" /> ID

Giuseppe Perale<fnref n="4" />,<fnref n="5" />,<fnref n="6" /> ID

Jonathan R. T. Jeffers<fnref n="11" /> ID

Filippo Rossi<fnref n="1" /> ID

Jonathan P. Wojciechowski<fnref n="7" />,<fnref n="8" />,<fnref n="9" /> ID

Molly M. Stevens<fnref n="7" />,<fnref n="8" />,<fnref n="9" /> ID

Correspondence

Håvard J. Haugen, Department of Biomaterials, Institute of Clinical Dentistry, University of Oslo, Geitmysrveien 71, 0455 Oslo, Norway.

Email: h.j.haugen@odont.uio.no

Funding information

Aker Scholarship; H2020 Marie Skłodowska-Curie Actions, Grant/Award Number: 801133; Norges Forskningsråd, Grant/Award Number: 327512; UK Research and Innovation, Grant/Award Number: MR/R015651/1

Abstract

Musculoskeletal defects are an enormous healthcare burden and source of pain and disability for individuals. With an aging population, the proportion of individuals living with these medical indications will increase. Simultaneously, there is pressure on healthcare providers to source efficient solutions, which are cheaper and less invasive than conventional technology. This has led to an increased research focus on hydrogels as highly biocompatible biomaterials that can be delivered through minimally invasive procedures. This review will discuss how hydrogels can be designed for clinical translation, particularly in the context of the new European Medical Device Regulation (MDR). We will then do a deep dive into the clinically used hydrogel solutions that have been commercially approved or have undergone clinical trials in Europe or the United States. We will discuss the therapeutic mechanism and limitations of these products. Due to the vast application areas of hydrogels, this work focuses only on treatments of cartilage, bone, and the nucleus pulposus. Lastly, the main steps toward clinical translation of hydrogels as medical devices are outlined. We suggest a framework for how academics can assist small and medium MedTech enterprises conducting the initial clinical investigation and post-market clinical follow-up required in the MDR. It is evident that the successful translation of hydrogels is governed by acquiring high-quality pre-clinical and clinical data confirming the device mechanism of action and safety.

KEYWORDS

bone regeneration, cartilage regeneration, clinical translation, hydrogels, medical devices, regenerative medicine, viscosupplementation

<footnote n="1">¹Department of Chemistry, Materials and Chemical Engineering “Giulio Natta”, Politecnico di Milano, Milano, Italy</footnote>

<footnote n="2">²Department of Biomaterials, Institute of Clinical Dentistry, University of Oslo, Oslo, Norway</footnote>

<footnote n="3">³Material Biomimetic AS, Oslo Science Park, Oslo, Norway</footnote>

<footnote n="4">⁴Industrie Biomediche Insubri SA, Mezzovico-Vira, Switzerland</footnote>

<footnote n="5">⁵Faculty of Biomedical Sciences, University of Southern Switzerland, Lugano, Switzerland</footnote>

<footnote n="6">⁶Ludwig Boltzmann Institute for Experimental and Clinical Traumatology, Vienna, Austria</footnote>

<footnote n="7">⁷Department of Materials, Imperial College London, London, UK</footnote>

<footnote n="8">⁸Department of Bioengineering, Imperial College London, London, UK</footnote>

<footnote n="9">⁹Institute of Biomedical Engineering, Imperial College London, London, UK</footnote>

<footnote n="10">¹⁰Hybrid Technology Hub, Centre of Excellence, Institute of Basic Medical Science, University of Oslo, Oslo, Norway</footnote>

<footnote n="11">¹¹Department of Mechanical Engineering, Imperial College London, London, UK</footnote>
# 1 | HYDROGELS AS MEDICAL DEVICES

Hydrogels represent a group of biomaterials consisting of water-swollen polymer or colloidal networks.¹ Hydrogels are viscoelastic materials that have attracted attention in regenerative medicine due to their ability to structurally mimic the extracellular matrix (ECM),² thereby creating a conducive environment for cell proliferation and tissue regeneration. The viscoelastic properties of hydrogels allow them to function as stem cell carriers or scaffolds for controlled drug release. Within the review by Correa and colleagues,³ these applications of hydrogels are discussed in the context of clinical translation. From a regulatory perspective, hydrogels can be considered medical devices if their therapeutic effect comes from their intrinsic structure, because their physical, chemical, or mechanical effects are the primary mechanism of action for their therapeutic function. To be classified as medical devices hydrogels cannot have any medicinal component, effect, or mechanism of action. For the case of medical devices, harmonized in the European market, the relevant regulation is EU 2017/745, which entered in force on May 26, 2021. Melvin and Torre⁴ have discussed the rationale behind the new regulation. Recently, Catoira et al.⁵ have discussed how the regulation affects the translation of hydrogels. Hence, this review will consider how hydrogels can be designed to satisfy these regulations. However, expanding cells or integrating cell-stimulating therapeutics into the medical device results in these hydrogel systems being regulated as medicinal products (drugs/biologics). Indeed, they are considered as drugs when their principal mode of action is pharmacological, metabolic, or immunological.⁶ The consequence of the medicinal regulation is that a more thorough investigation of the biocompatibility and therapeutic effect is required before such solutions can be approved for clinical application. This increases the translational barriers and the time before patients can benefit from the treatment. Therefore, it is attractive to translate hydrogels solutions as medical devices such that the therapy can reach the clinic earlier and is more affordable. A detailed discussion of the classification can be found in Section 4.

Particularly in applications for musculoskeletal disorders, there is an unmet need for minimally invasive therapies, where the use of injectable hydrogels has tremendous potential. The demand is driven by an aging population that gives two unique opportunities: (1) an increasing number of patients outlive the longevity of permanent medical devices; thus, hydrogel therapies can be used to delay permanent implantation, (2) minimally invasive therapies give a treatment opportunity for the growing population group that would otherwise not survive the trauma induced by conventional surgeries.⁷ Examples of these type of devices are represented by hydrogels for joint lubrication,⁸,⁹ injectable scaffolds for guided bone¹⁰ or cartilage regeneration,¹¹ or nucleus pulposus (NP) replacements.¹² These are widely different applications for diverse tissues with different loading modes and levels. Consequently, a one-fit-all hydrogel is an unlikely strategy, and thought should be put into the clinical requirements of the material when designing the hydrogel.

# 2 | HYDROGEL DESIGN

Hydrogels are an extensively investigated class of biomaterials, and an increasing number of products have reached the clinic. In the following section, we will go through the design steps of the hydrogels and discuss what considerations need to be taken to improve the likelihood of clinical translation and to comply with the European Medical Device Regulations (MDRs). The design process is summarized in Figure 1.

## 2.1 | Material selection

The first step in the design process is to select a suitable polymer to form the hydrogel. There is a larger group of naturally derived polymers such as collagen,¹³ hyaluronic acid (HA),¹⁴ chitosan,¹⁵ cellulose,¹⁶ and alginate.¹⁷ Although not exclusively, plant-based polymers tend to be composed of saccharides, such as cellulose, and animal-based polymers tend to compose of protein, for example, collagen.¹⁸ These have been attractive as their natural origin makes them favorably biocompatible and biodegradable but can introduce issues such as immunogenicity and limited mechanical properties.¹⁹ From a translational perspective, these are limited by high cost and batch-to-batch variability.⁵,¹⁹,²⁰ Alternatively, synthetic polymers such as poly(ethylene glycol) (PEG),²¹ poly(vinyl alcohol),²² poly(acrylic acid),²³ and poly(acrylamide)²⁴ can be used. Synthetic polymers are industrially more used as they are more favorable from both cost and regulatory perspective, the two being also connected. Synthetic polymers can be produced in more robustly repeatable manners and more efficiently with respect to naturally derived ones, making them readily scalable.²⁵ Synthesis is typically a more straightforward production process and ensures controlled environmental factors thus limiting the risk of contamination. Synthetic polymers are favorable versus naturally derived raw materials as they allow for improved traceability and higher degree of availability which finally reduces the cost.²⁵,²⁶ However, their clinical adoption has been limited, and those that exist usually provide a mechanical mechanism of action, for example, PEG-hydrogel as a spacer between prostate and rectum to protect the rectum during radiotherapy.²⁷ For the regenerative market, the translation is insignificant, which has been attributed to their low biocompatibility.²⁸ The low biocompatibility is likely related to lack of cell-specific bioactivity, including cell adhesive and migratory cues, and cell-mediated material degradation.²⁹ This highlights how biocompatibility is vital for the success of any hydrogel, and the biological response should be central to the choice of the polymer for the hydrogel. Implanted materials can either integrate physiologically, leading to minimal or no scaring, or the material can induce chronic inflammation and a foreign body response.³⁰ After injection, the material must provide appropriate biochemical and biophysical signals to recruit host cells that will eventually produce new native tissue.³¹ Immune cells also play a key role in the signaling cascade leading to tissue regeneration, and appropriate engineering of the local immune response can boost the tissue regeneration.³² For instance,
## 1 | HYDROGELS AS MEDICAL DEVICES

Hydrogels represent a group of biomaterials consisting of water-swollen polymer or colloidal networks.¹ Hydrogels are viscoelastic materials that have attracted attention in regenerative medicine due to their ability to structurally mimic the extracellular matrix (ECM),² thereby creating a conducive environment for cell proliferation and tissue regeneration. The viscoelastic properties of hydrogels allow them to function as stem cell carriers or scaffolds for controlled drug release. Within the review by Correa and colleagues,³ these applications of hydrogels are discussed in the context of clinical translation. From a regulatory perspective, hydrogels can be considered medical devices if their therapeutic effect comes from their intrinsic structure, because their physical, chemical, or mechanical effects are the primary mechanism of action for their therapeutic function. To be classified as medical devices hydrogels cannot have any medicinal component, effect, or mechanism of action. For the case of medical devices, harmonized in the European market, the relevant regulation is EU 2017/745, which entered in force on May 26, 2021. Melvin and Torre⁴ have discussed the rationale behind the new regulation. Recently, Catoira et al.⁵ have discussed how the regulation affects the translation of hydrogels. Hence, this review will consider how hydrogels can be designed to satisfy these regulations. However, expanding cells or integrating cell-stimulating therapeutics into the medical device results in these hydrogel systems being regulated as medicinal products (drugs/biologics). Indeed, they are considered as drugs when their principal mode of action is pharmacological, metabolic, or immunological.⁶ The consequence of the medicinal regulation is that a more thorough investigation of the biocompatibility and therapeutic effect is required before such solutions can be approved for clinical application. This increases the translational barriers and the time before patients can benefit from the treatment. Therefore, it is attractive to translate hydrogels solutions as medical devices such that the therapy can reach the clinic earlier and is more affordable. A detailed discussion of the classification can be found in Section 4.

Particularly in applications for musculoskeletal disorders, there is an unmet need for minimally invasive therapies, where the use of injectable hydrogels has tremendous potential. The demand is driven by an aging population that gives two unique opportunities: (1) an increasing number of patients outlive the longevity of permanent medical devices; thus, hydrogel therapies can be used to delay permanent implantation, (2) minimally invasive therapies give a treatment opportunity for the growing population group that would otherwise not survive the trauma induced by conventional surgeries.⁷ Examples of these type of devices are represented by hydrogels for joint lubrication,⁸,⁹ injectable scaffolds for guided bone¹⁰ or cartilage regeneration,¹¹ or nucleus pulposus (NP) replacements.¹² These are widely different applications for diverse tissues with different loading modes and levels. Consequently, a one-fit-all hydrogel is an unlikely strategy, and thought should be put into the clinical requirements of the material when designing the hydrogel.

## 2 | HYDROGEL DESIGN

Hydrogels are an extensively investigated class of biomaterials, and an increasing number of products have reached the clinic. In the following section, we will go through the design steps of the hydrogels and discuss what considerations need to be taken to improve the likelihood of clinical translation and to comply with the European Medical Device Regulations (MDRs). The design process is summarized in Figure 1.

### 2.1 | Material selection

The first step in the design process is to select a suitable polymer to form the hydrogel. There is a larger group of naturally derived polymers such as collagen,¹³ hyaluronic acid (HA),¹⁴ chitosan,¹⁵ cellulose,¹⁶ and alginate.¹⁷ Although not exclusively, plant-based polymers tend to be composed of saccharides, such as cellulose, and animal-based polymers tend to compose of protein, for example, collagen.¹⁸ These have been attractive as their natural origin makes them favorably biocompatible and biodegradable but can introduce issues such as immunogenicity and limited mechanical properties.¹⁹ From a translational perspective, these are limited by high cost and batch-to-batch variability.⁵,¹⁹,²⁰ Alternatively, synthetic polymers such as poly(ethylene glycol) (PEG),²¹ poly(vinyl alcohol),²² poly(acrylic acid),²³ and poly(acrylamide)²⁴ can be used. Synthetic polymers are industrially more used as they are more favorable from both cost and regulatory perspective, the two being also connected. Synthetic polymers can be produced in more robustly repeatable manners and more efficiently with respect to naturally derived ones, making them readily scalable.²⁵ Synthesis is typically a more straightforward production process and ensures controlled environmental factors thus limiting the risk of contamination. Synthetic polymers are favorable versus naturally derived raw materials as they allow for improved traceability and higher degree of availability which finally reduces the cost.²⁵,²⁶ However, their clinical adoption has been limited, and those that exist usually provide a mechanical mechanism of action, for example, PEG-hydrogel as a spacer between prostate and rectum to protect the rectum during radiotherapy.²⁷ For the regenerative market, the translation is insignificant, which has been attributed to their low biocompatibility.²⁸ The low biocompatibility is likely related to lack of cell-specific bioactivity, including cell adhesive and migratory cues, and cell-mediated material degradation.²⁹ This highlights how biocompatibility is vital for the success of any hydrogel, and the biological response should be central to the choice of the polymer for the hydrogel. Implanted materials can either integrate physiologically, leading to minimal or no scaring, or the material can induce chronic inflammation and a foreign body response.³⁰ After injection, the material must provide appropriate biochemical and biophysical signals to recruit host cells that will eventually produce new native tissue.³¹ Immune cells also play a key role in the signaling cascade leading to tissue regeneration, and appropriate engineering of the local immune response can boost the tissue regeneration.³² For instance,
<figure><img src="image_3.png" /><figcaption>**FIGURE 1** Schematic of the design process of hydrogels as medical devices for musculoskeletal application. The process consists of three design blocks, first, the hydrogel is developed, then any particles or other composite inclusions are added before the delivery strategy is chosen</figcaption></figure>

monocytes and macrophages releasing cytokines including BMP-2, BP-4, and TGF-β1 support osteoblast differentiation and proliferation.33 The current gold standard for understanding biocompatibility remains clinical trials, but essential information can also be derived from well-designed pre-clinical trials. When selecting the polymer, we have two conflicting interests; from a biological perspective, natural biopolymers are favorable due to higher biocompatibility, meanwhile synthetic polymers have more controllable properties, including swelling, degradation, phase transitions, and mechanical properties.34 Additionally, synthetic polymers are more favorable from a regulatory and financial perspective. To balance these interests co-gel solutions such as PEG-HA35 and gelatin methacrylate-PEG diacrylate36 are promising strategies at combining features from both groups of polymers. Moreover, synthetic polymers can be functionalized with proteins and peptides to improve cell attachment and proliferation. For instance, the inclusion of RGD-peptides in PEG gels has demonstrated these capabilities on multiple cell types.37,38

## 2.2 | Crosslinking/gelation

The next step is to form the gel-network by crosslinking the polymer chains. There are two options here, and the polymer can be physically

or chemically crosslinked. Physical crosslinking is a reversible process where weak non-covalent interactions (e.g., van der Waals, hydrogen bonding, electrostatic interactions) keep the network stable. The advantage here is that the gel can be formed without using a crosslinking agent and the gel is easier to mold into the defect geometry.39 Alternatively, chemical crosslinking (covalent bonds) can be used. The covalent bonds tend to convey to the gel's improved mechanical properties and higher stability.40 The gel stability is a vital matrix design as long degradation times and the inability to be remodeled by the cells will hamper tissue growth. In contrast, a fast degradation time will leave an unfilled void after the gel degradation.41 The degree of crosslinking, meaning the number of bonds that interconnect the polymers to each other, is an important parameter for the material properties. With a higher degree of crosslinking, we can expect a higher viscosity, stiffness, and longer degradation time.42,43 It has early been established that with increased degree of crosslinking, the gel's ability to swell decreases.44 The equilibrium degree of swelling affects a series of properties such as solute diffusion coefficient, mechanical properties, and the mobility of therapeutic agents.45 From a translational perspective, chemical crosslinking means introducing new chemicals and at least one more chemical reaction. It must be proven that the biomaterial remains biocompatible and that there is not an increase of leachables such as unreacted crosslinker. Dialysis tends to be an efficient method for removing such
<figure><img src="image_3.png" /><figcaption>**FIGURE 1** Schematic of the design process of hydrogels as medical devices for musculoskeletal application. The process consists of three design blocks, first, the hydrogel is developed, then any particles or other composite inclusions are added before the delivery strategy is chosen</figcaption></figure>

monocytes and macrophages releasing cytokines including BMP-2, BP-4, and TGF-β1 support osteoblast differentiation and proliferation.33 The current gold standard for understanding biocompatibility remains clinical trials, but essential information can also be derived from well-designed pre-clinical trials. When selecting the polymer, we have two conflicting interests; from a biological perspective, natural biopolymers are favorable due to higher biocompatibility, meanwhile synthetic polymers have more controllable properties, including swelling, degradation, phase transitions, and mechanical properties.34 Additionally, synthetic polymers are more favorable from a regulatory and financial perspective. To balance these interests co-gel solutions such as PEG-HA35 and gelatin methacrylate-PEG diacrylate36 are promising strategies at combining features from both groups of polymers. Moreover, synthetic polymers can be functionalized with proteins and peptides to improve cell attachment and proliferation. For instance, the inclusion of RGD-peptides in PEG gels has demonstrated these capabilities on multiple cell types.37,38

## 2.2 | Crosslinking/gelation

The next step is to form the gel-network by crosslinking the polymer chains. There are two options here, and the polymer can be physically

or chemically crosslinked. Physical crosslinking is a reversible process where weak non-covalent interactions (e.g., van der Waals, hydrogen bonding, electrostatic interactions) keep the network stable. The advantage here is that the gel can be formed without using a crosslinking agent and the gel is easier to mold into the defect geometry.39 Alternatively, chemical crosslinking (covalent bonds) can be used. The covalent bonds tend to convey to the gel's improved mechanical properties and higher stability.40 The gel stability is a vital matrix design as long degradation times and the inability to be remodeled by the cells will hamper tissue growth. In contrast, a fast degradation time will leave an unfilled void after the gel degradation.41 The degree of crosslinking, meaning the number of bonds that interconnect the polymers to each other, is an important parameter for the material properties. With a higher degree of crosslinking, we can expect a higher viscosity, stiffness, and longer degradation time.42,43 It has early been established that with increased degree of crosslinking, the gel's ability to swell decreases.44 The equilibrium degree of swelling affects a series of properties such as solute diffusion coefficient, mechanical properties, and the mobility of therapeutic agents.45 From a translational perspective, chemical crosslinking means introducing new chemicals and at least one more chemical reaction. It must be proven that the biomaterial remains biocompatible and that there is not an increase of leachables such as unreacted crosslinker. Dialysis tends to be an efficient method for removing such
impurities, but the removal and biocompatibility must be proven. This
is further discussed in Section 5.

## 2.3 | Composite design

Like with other biomaterial types, composite materials can be formed using hydrogels. This is a favorable strategy as the final material will have inherent properties from base materials in addition to the properties derived from the interaction between material components. In terms of hydrogels, this could be the introduction of fibers, for instance, to improve mechanical properties or guide tissue growth, or particles, for example, a ceramic phase which can boost bone regeneration. Li and co-workers⁴⁶ combined a hydrogel of thiolated HA and polyethylene glycol diacrylate (PEGDA) gel covalently crosslinked to fragmented, electrospun polycaprolactone fibers. The fibers gave improved mechanical properties compared to the HA-PEG gel alone, thereby they could mimic the mechanical properties of native fat tissue. Moreover, their in vivo trials with subcutaneous injection of the material in a rat and a rabbit model suggest improved macrophage polarization toward a pro-regenerative phenotype and enhanced angiogenesis.

The inclusion of an inorganic phase in the polymeric hydrogel material has been a popular strategy for bone regeneration. Chahal and colleagues⁴⁷ developed a PEG hydrogel with amorphous calcium phosphate particles. They demonstrated that the particles both gave a higher stiffness, and slowly released calcium and zinc ions into the solution, creating conducive properties for bone regeneration. Although they observed a qualitative increase in gel mineralization, they could not demonstrate statistical significance. Furthermore, the human mesenchymal stem cells (MSCs) they used were unable to attach to the gel before they functionalized the PEG with RGD tripeptide motifs. This highlights the importance of choosing a polymer with high bioactivity to succeed clinically and demonstrates one of the shortfalls of most fully synthetic systems. Semi-synthetic systems, however, are promising as they allow for tunable properties such as gelation mechanism and adhesion to tissues. Researchers from the Langer lab⁴⁸ developed a cellulose hydrogel with PEG-block-poly (lactic acid) nanoparticles as non-covalent crosslinking nodes that gave the gel shear-thinning and self-healing properties. In vivo in mice (subcutaneously in the back) they demonstrated biocompatibility with a mild neutrophil-induced inflammation at day 3 and clearance by macrophages from day 7. A consistent release pattern was observed when particles were loaded with model dual-hydrophobic/hydrophilic drugs. Wang and colleagues⁴⁹ formed a mechanically strong, transparent, and self-healing hydrogel by coating clay nanosheets with sodium polyacrylate and physically crosslinking it with dendritic G2 binder.

Any inclusion will make it, from a regulatory perspective, a completely new biomaterial. Therefore, it will require the standard omni-comprehensive testing due for any novel formulation. This includes the application of ISO 10993 family of standards that encompass biocompatibility testing up to clinical studies, and for biodegradable biomaterials the documentation requirements that degradation products do not accumulate in any body organs.⁵⁰

## 2.4 | Implantation method

Although the implantation method of the hydrogel is not directly a design variable affecting the gel properties, the hydrogel should be designed with implantation feasibility in mind. There are primarily two strategies of implantation in current use. The traditional is surgical incision implantation, where a surgeon cuts a flap through the patient's dermis and physically places the implant at the desired location. The advantage of this intervention is that the gel can be pre-shaped prior to the surgery and have higher mechanical stiffness. The disadvantage is that the incision surgery gives a longer hospitalization time, longer recovery time, increased postoperative pain,⁵¹ and higher risk of bacterial infections.⁵² Therefore, injectable solutions are attractive minimally invasive strategies that give less trauma,⁵³ less blood loss, shorter surgeries, and rapid recovery.⁵⁴ This brings its own technical challenges, as the gel must have low enough viscosity to be injectable through a needle or arthroscopic instruments. To have adequate viscosity during injection, it might be favorable to use a low degree of crosslinking,⁴³ a physically crosslinked gel exhibiting shear-thinning properties,⁴⁸,⁵⁵ or utilize in situ crosslinking of the hydrogel using methods such as click-chemistry,⁵⁶ ultrasound,⁵⁷ and photo-initiated crosslinking.¹¹,¹² For in situ crosslinking, it is imperative to ensure that there are no adverse chemical reactions between the material and the surrounding biological tissue. For instance, thiol groups are naturally occurring in the body, so if a thiol-based Michael addition strategy is used for gelation, there is a risk of undesired cross-reactivity, oxidation, or metabolism.⁴⁰ This has inspired the focus on bioorthogonal chemistry, a class of high-yielding reactions based on selective transformation not commonly found in biology,⁵⁸

An innovative solution for injection of a hydrogel therapy is the Flowbone® solution developed by researchers at EPFL in Switzerland. They have developed a biphasic gel solution for bone regeneration where the first phase consists of covalently crosslinked HA with hydroxyapatite particles incorporated, that is carried in a second aqueous phase comprising more hydroxyapatite particles.⁵⁹ The biphasic system allows a low viscosity and thereby injectability. This solution also allows for the loading drugs such as bisphosphonates,⁶⁰ which is now under investigation in pre-clinical trials.

Other solutions chose a tactic where the crosslinking occurs in situ, such as Regentis Biomaterial's GelrinC®, which is discussed below. The in situ strategy allows for a low viscosity during injection, while the high viscosity and mechanical properties are obtained after injection.

## 3 | APPLICATIONS

A series of hydrogel-based products have been approved for clinical use in the EU and the United States, particularly for viscosupplementation (VS) in joints for osteoarthritis (OA). Furthermore, regenerative gels are now emerging that in addition to providing temporary pain relief and functional improvement, attempt to regrow or support the regrowth of the tissue for a longer-lasting therapeutic
impurities, but the removal and biocompatibility must be proven. This
is further discussed in Section 5.

## 2.3 | Composite design

Like with other biomaterial types, composite materials can be formed using hydrogels. This is a favorable strategy as the final material will have inherent properties from base materials in addition to the properties derived from the interaction between material components. In terms of hydrogels, this could be the introduction of fibers, for instance, to improve mechanical properties or guide tissue growth, or particles, for example, a ceramic phase which can boost bone regeneration. Li and co-workers⁴⁶ combined a hydrogel of thiolated HA and polyethylene glycol diacrylate (PEGDA) gel covalently crosslinked to fragmented, electrospun polycaprolactone fibers. The fibers gave improved mechanical properties compared to the HA-PEG gel alone, thereby they could mimic the mechanical properties of native fat tissue. Moreover, their in vivo trials with subcutaneous injection of the material in a rat and a rabbit model suggest improved macrophage polarization toward a pro-regenerative phenotype and enhanced angiogenesis.

The inclusion of an inorganic phase in the polymeric hydrogel material has been a popular strategy for bone regeneration. Chahal and colleagues⁴⁷ developed a PEG hydrogel with amorphous calcium phosphate particles. They demonstrated that the particles both gave a higher stiffness, and slowly released calcium and zinc ions into the solution, creating conducive properties for bone regeneration. Although they observed a qualitative increase in gel mineralization, they could not demonstrate statistical significance. Furthermore, the human mesenchymal stem cells (MSCs) they used were unable to attach to the gel before they functionalized the PEG with RGD tripeptide motifs. This highlights the importance of choosing a polymer with high bioactivity to succeed clinically and demonstrates one of the shortfalls of most fully synthetic systems. Semi-synthetic systems, however, are promising as they allow for tunable properties such as gelation mechanism and adhesion to tissues. Researchers from the Langer lab⁴⁸ developed a cellulose hydrogel with PEG-block-poly (lactic acid) nanoparticles as non-covalent crosslinking nodes that gave the gel shear-thinning and self-healing properties. In vivo in mice (subcutaneously in the back) they demonstrated biocompatibility with a mild neutrophil-induced inflammation at day 3 and clearance by macrophages from day 7. A consistent release pattern was observed when particles were loaded with model dual-hydrophobic/hydrophilic drugs. Wang and colleagues⁴⁹ formed a mechanically strong, transparent, and self-healing hydrogel by coating clay nanosheets with sodium polyacrylate and physically crosslinking it with dendritic G2 binder.

Any inclusion will make it, from a regulatory perspective, a completely new biomaterial. Therefore, it will require the standard omni-comprehensive testing due for any novel formulation. This includes the application of ISO 10993 family of standards that encompass biocompatibility testing up to clinical studies, and for biodegradable biomaterials the documentation requirements that degradation products do not accumulate in any body organs.⁵⁰

## 2.4 | Implantation method

Although the implantation method of the hydrogel is not directly a design variable affecting the gel properties, the hydrogel should be designed with implantation feasibility in mind. There are primarily two strategies of implantation in current use. The traditional is surgical incision implantation, where a surgeon cuts a flap through the patient's dermis and physically places the implant at the desired location. The advantage of this intervention is that the gel can be pre-shaped prior to the surgery and have higher mechanical stiffness. The disadvantage is that the incision surgery gives a longer hospitalization time, longer recovery time, increased postoperative pain,⁵¹ and higher risk of bacterial infections.⁵² Therefore, injectable solutions are attractive minimally invasive strategies that give less trauma,⁵³ less blood loss, shorter surgeries, and rapid recovery.⁵⁴ This brings its own technical challenges, as the gel must have low enough viscosity to be injectable through a needle or arthroscopic instruments. To have adequate viscosity during injection, it might be favorable to use a low degree of crosslinking,⁴³ a physically crosslinked gel exhibiting shear-thinning properties,⁴⁸,⁵⁵ or utilize in situ crosslinking of the hydrogel using methods such as click-chemistry,⁵⁶ ultrasound,⁵⁷ and photo-initiated crosslinking.¹¹,¹² For in situ crosslinking, it is imperative to ensure that there are no adverse chemical reactions between the material and the surrounding biological tissue. For instance, thiol groups are naturally occurring in the body, so if a thiol-based Michael addition strategy is used for gelation, there is a risk of undesired cross-reactivity, oxidation, or metabolism.⁴⁰ This has inspired the focus on bioorthogonal chemistry, a class of high-yielding reactions based on selective transformation not commonly found in biology,⁵⁸

An innovative solution for injection of a hydrogel therapy is the Flowbone® solution developed by researchers at EPFL in Switzerland. They have developed a biphasic gel solution for bone regeneration where the first phase consists of covalently crosslinked HA with hydroxyapatite particles incorporated, that is carried in a second aqueous phase comprising more hydroxyapatite particles.⁵⁹ The biphasic system allows a low viscosity and thereby injectability. This solution also allows for the loading drugs such as bisphosphonates,⁶⁰ which is now under investigation in pre-clinical trials.

Other solutions chose a tactic where the crosslinking occurs in situ, such as Regentis Biomaterial's GelrinC®, which is discussed below. The in situ strategy allows for a low viscosity during injection, while the high viscosity and mechanical properties are obtained after injection.

## 3 | APPLICATIONS

A series of hydrogel-based products have been approved for clinical use in the EU and the United States, particularly for viscosupplementation (VS) in joints for osteoarthritis (OA). Furthermore, regenerative gels are now emerging that in addition to providing temporary pain relief and functional improvement, attempt to regrow or support the regrowth of the tissue for a longer-lasting therapeutic
<table><caption>TABLE 1 Clinically available injectable hydrogel products for treatments of musculoskeletal indications</caption><thead><tr><th>Indication and treatment mode</th><th>Product (manufacturer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>CE/FDA approvals</th></tr></thead><tbody><tr><td>Osteoarthritis</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>VS</td><td>Gel-One®61,62 (Zimmer Biomet)</td><td>Cinnamic acid functionalized HA crosslinked with UV light</td><td>Single 3 mL injection</td><td>Pain relief up to 26 weeks</td><td>FDA</td></tr><tr><td></td><td>Orthovisc®63 (Anika Therapeutics)</td><td>SHA (1.0–2.9 MDa) in saline water</td><td>Three separate injections of 30 mg in 2 mL solution</td><td>Pain relief up to 6 months</td><td>FDA</td></tr><tr><td></td><td>Monovisc®64 (Anika Therapeutics)</td><td>High MW HA, lightly crosslinked with biscarbodiimide</td><td>One 4 mL injection</td><td>6 months pain relief</td><td>FDA</td></tr><tr><td></td><td>Hymovis®65 (Fidia Farmaceutici)</td><td>HA 500–730 kDa, functionalized with 2–3% hexadecylamine</td><td>Two times 3 mL at a week interval</td><td>Lubrication and pain relief up to 12 months</td><td>FDA</td></tr><tr><td></td><td>Arthrosamid®66,67 (Conura International)</td><td>2.5% polyacrylamide in sterile water–non-degradable</td><td>Single session injection of 6 × 1 mL through 21G cannula (syringes replaced using luer-lock system)</td><td>Long-lasting pain relief (52 weeks proven)<sup>a</sup></td><td>CE &amp; Clinical Trials (US)</td></tr><tr><td>Regeneration</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td></td><td>BST-Car-Gel®68,69 (Smith &amp; Nephew)</td><td>Chitosan dissolved in saline water and autologous blood</td><td>Mini-arthrotomy or arthroscopy in combination with bone marrow stimulations such as microfracture</td><td>Superior hyaline cartilage regeneration compared to microfracture alone</td><td>CE</td></tr><tr><td></td><td>RegenoGel®70 (ProCore)</td><td>HA (1.6 MDa) conjugated to purified platelet-rich plasma-derived fibrinogen</td><td>Synovial fluid is first removed through a 21G needle, before 4 mL of gel is injected. Two administrations 3 months apart</td><td>Pain relief and cartilage regeneration</td><td>Clinical trials (EU and US)</td></tr><tr><td>Bone defects<sup>b</sup></td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Regeneration</td><td>GelrinC®71–73 (Regentis Biomat.)</td><td>PEGDA with denatured fibrinogen</td><td>Injected after microfracture, crosslinked using UVA light</td><td>Degrades over 6–12 months while being replaced by regenerated cartilage</td><td>Clinical trials (EU and US)</td></tr><tr><td>Regeneration</td><td>Emdogain®74,75 (Straumann)</td><td>Porcine EMD in propylene glycol alginate gel [30 mg/m]</td><td>Flap incision or flapless injection in dental application</td><td>Regenerates periodontal tissue (cementum, periodontal ligament, bone)</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Perioglas®76 (NovaBone)</td><td>Calcium phosphosilicate particles + a PEG and glycerine gel-like binder</td><td>Either as a moldable putty or through syringe injection</td><td>Dental bone regeneration</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Actifuse®77,78 (Baxter)</td><td>Phase-pure silicon-substituted calcium phosphate in poloxamer 407</td><td>Injectable through syringe</td><td>Bone void filler in spinal and orthopedic application</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Dynagraft III®79 (Integra)</td><td>DBM in poloxamer carrier</td><td>Injectable through syringe or delivered as a putty</td><td>Bone void filler</td><td>FDA</td></tr><tr><td></td><td>AlphaGRAFT® (Alphatech)</td><td>DBM in poloxamer carrier</td><td>Extruded through syringe</td><td>Bone void filler</td><td>FDA</td></tr><tr><td></td><td>AlloFuse®79 (AlloSource)</td><td>29% allographic DBM in polyethylene oxide polypropylene oxide block copolymer</td><td>Mixed with autologous bone for spinal fusion or injected in trauma cases</td><td>Void filler, graft extender</td><td>FDA</td></tr><tr><td></td><td>Optium®79 (LifeNet Health)</td><td>Allographic DBM in glycerol</td><td>Allographic DBM in glycerol</td><td>Bone graft extender and void filler</td><td>FDA</td></tr></tbody><tfoot><tr><td colspan="6" style="text-align: right;">(Continues)</td></tr></tfoot></table>
<table><caption>TABLE 1 Clinically available injectable hydrogel products for treatments of musculoskeletal indications</caption><thead><tr><th>Indication and treatment mode</th><th>Product (manufacturer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>CE/FDA approvals</th></tr></thead><tbody><tr><td>Osteoarthritis</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>VS</td><td>Gel-One®61,62 (Zimmer Biomet)</td><td>Cinnamic acid functionalized HA crosslinked with UV light</td><td>Single 3 mL injection</td><td>Pain relief up to 26 weeks</td><td>FDA</td></tr><tr><td></td><td>Orthovisc®63 (Anika Therapeutics)</td><td>SHA (1.0–2.9 MDa) in saline water</td><td>Three separate injections of 30 mg in 2 mL solution</td><td>Pain relief up to 6 months</td><td>FDA</td></tr><tr><td></td><td>Monovisc®64 (Anika Therapeutics)</td><td>High MW HA, lightly crosslinked with biscarbodiimide</td><td>One 4 mL injection</td><td>6 months pain relief</td><td>FDA</td></tr><tr><td></td><td>Hymovis®65 (Fidia Farmaceutici)</td><td>HA 500–730 kDa, functionalized with 2–3% hexadecylamine</td><td>Two times 3 mL at a week interval</td><td>Lubrication and pain relief up to 12 months</td><td>FDA</td></tr><tr><td></td><td>Arthrosamid®66,67 (Contra International)</td><td>2.5% polyacrylamide in sterile water–non-degradable</td><td>Single session injection of 6 × 1 mL through 21G cannula (syringes replaced using luer-lock system)</td><td>Long-lasting pain relief (52 weeks proven)<sup>a</sup></td><td>CE &amp; Clinical Trials (US)</td></tr><tr><td>Regeneration</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td></td><td>BST-Car-Gel®68,69 (Smith &amp; Nephew)</td><td>Chitosan dissolved in saline water and autologous blood</td><td>Mini-arthrotomy or arthroscopy in combination with bone marrow stimulations such as microfracture</td><td>Superior hyaline cartilage regeneration compared to microfracture alone</td><td>CE</td></tr><tr><td></td><td>RegenoGel®70 (ProCore)</td><td>HA (1.6 MDa) conjugated to purified platelet-rich plasma-derived fibrinogen</td><td>Synovial fluid is first removed through a 21G needle, before 4 mL of gel is injected. Two administrations 3 months apart</td><td>Pain relief and cartilage regeneration</td><td>Clinical trials (EU and US)</td></tr><tr><td>Bone defects<sup>b</sup></td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Regeneration</td><td>GelrinC®71–73 (Regentis Biomat.)</td><td>PEGDA with denatured fibrinogen</td><td>Injected after microfracture, crosslinked using UVA light</td><td>Degrades over 6–12 months while being replaced by regenerated cartilage</td><td>Clinical trials (EU and US)</td></tr><tr><td>Regeneration</td><td>Emdogain®74,75 (Straumann)</td><td>Porcine EMD in propylene glycol alginate gel [30 mg/m]</td><td>Flap incision or flapless injection in dental application</td><td>Regenerates periodontal tissue (cementum, periodontal ligament, bone)</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Perioglas®76 (NovaBone)</td><td>Calcium phosphosilicate particles + a PEG and glycerine gel-like binder</td><td>Either as a moldable putty or through syringe injection</td><td>Dental bone regeneration</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Actifuse®77,78 (Baxter)</td><td>Phase-pure silicon-substituted calcium phosphate in poloxamer 407</td><td>Injectable through syringe</td><td>Bone void filler in spinal and orthopedic application</td><td>CE &amp; FDA</td></tr><tr><td></td><td>Dynagraft III®79 (Integra)</td><td>DBM in poloxamer carrier</td><td>Injectable through syringe or delivered as a putty</td><td>Bone void filler</td><td>FDA</td></tr><tr><td></td><td>AlphaGRAFT® (Alphatech)</td><td>DBM in poloxamer carrier</td><td>Extruded through syringe</td><td>Bone void filler</td><td>FDA</td></tr><tr><td></td><td>AlloFuse®79 (AlloSource)</td><td>29% allographic DBM in polyethylene oxide polypropylene oxide block copolymer</td><td>Mixed with autologous bone for spinal fusion or injected in trauma cases</td><td>Void filler, graft extender</td><td>FDA</td></tr><tr><td></td><td>Optium®79 (LifeNet Health)</td><td>Allographic DBM in glycerol</td><td>Allographic DBM in glycerol</td><td>Bone graft extender and void filler</td><td>FDA</td></tr></tbody><tfoot><tr><td colspan="6" style="text-align: right;">(Continues)</td></tr></tfoot></table>
<table><caption>TABLE 1 (Continued)</caption><thead><tr><th>Indication and treatment mode</th><th>Product (manufacturer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>CE/FDA approvals</th></tr></thead><tbody><tr><td></td><td>Grafton DBM® gel<sup>80,81</sup> (Medtronic)</td><td>Allographic DBM in glycerol</td><td>Mixed with autologous bone for spinal fusion or injected in trauma cases</td><td>Bone graft extender and void filler</td><td>FDA</td></tr><tr><td></td><td>Tactoset<sup>82</sup> (Anika Therapeutics)</td><td>HA carrier with calcium phosphate</td><td>The HA and CaP are mixed then injected. Hardens within 15-20 min</td><td>Bone void filler for orthopedic application</td><td>FDA</td></tr><tr><td></td><td>Kinex® Bioactive Gel<sup>83</sup> (Globus Medical)</td><td>Bioglass, collagen and HA</td><td>Injectable solution</td><td>Bone void filler</td><td>FDA</td></tr><tr><td colspan="6"><strong>DDD</strong></td></tr><tr><td rowspan="4">Nucleus pulposus replacement</td><td>GelStix<sup>84</sup> (Replication Med.)</td><td>Polyacrylonitrile</td><td>Injected through a 22G needle and swells in situ</td><td>Pain relief from 1 month after surgery for at least 12 months</td><td>CE</td></tr><tr><td>HYADD4-G<sup>85,86</sup> (Fidia Farmaceutici)</td><td>HA 500-730 kDa, functionalized with 2%-3% hexadecylamine</td><td>1.5 mL [8 mg/mL] intradiscal injections guided by X-ray</td><td>Statistically significant pain relief up to 24 weeks</td><td>Clinical Trials</td></tr><tr><td>BioDisc<sup>87,88</sup> (CryoLife)</td><td>Albumin + glutaraldehyde hydrogel</td><td>Crosslink in situ within 2 min</td><td>Reduction in pain after 6 months</td><td>Unknown<sup>c</sup></td></tr><tr><td>NuCore<sup>89,90</sup> (Spine Wave)</td><td>Block polymers of silk and elastin crosslinked in situ with diisocyanate</td><td>Injected with a sealed vented needle to recover disc height (0.3-1.9 mL)</td><td>Reduction in back and leg pain, regained disc height</td><td>Unknown<sup>c</sup></td></tr></tbody></table>

Note: The list of products for viscosupplementation (VS) is non-exhaustive due to many products on the market. Instead, representative products for different materials and crosslinking mechanisms are presented. The other types of products are exhaustive to the authors' best knowledge.

Abbreviations: CaP, calcium phosphate; DBM, demineralized bone matrix; DDD, degenerative disc disease; EMD, enamel matrix derivatives; FDA, Food and Drug Administration; HA, hyaluronic acid; PBS, phosphate-buffered saline; SHA, sodium hyaluronate; VS, viscosupplementation.

<footnote><sup>a</sup>Clinical trial has a 5-year follow-up period.</footnote>

<footnote><sup>b</sup>Bone defects = trauma, oncology, craniofacial.</footnote>

<footnote><sup>c</sup>Seems to be discontinued.</footnote>
<table><caption>TABLE 1 (Continued)</caption><thead><tr><th>Indication and treatment mode</th><th>Product (manufacturer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>CE/FDA approvals</th></tr></thead><tbody><tr><td></td><td>Grafton DBM® gel<sup>80,81</sup> (Medtronic)</td><td>Allographic DBM in glycerol</td><td>Mixed with autologous bone for spinal fusion or injected in trauma cases</td><td>Bone graft extender and void filler</td><td>FDA</td></tr><tr><td></td><td>Tactoset<sup>82</sup> (Anika Therapeutics)</td><td>HA carrier with calcium phosphate</td><td>The HA and CaP are mixed then injected. Hardens within 15-20 min</td><td>Bone void filler for orthopedic application</td><td>FDA</td></tr><tr><td></td><td>Kinex® Bioactive Gel<sup>83</sup> (Globus Medical)</td><td>Bioglass, collagen and HA</td><td>Injectable solution</td><td>Bone void filler</td><td>FDA</td></tr><tr><td colspan="6"><strong>DDD</strong></td></tr><tr><td rowspan="5">Nucleus pulposus replacement</td><td>GelStix<sup>84</sup> (Replication Med.)</td><td>Polyacrylonitrile</td><td>Injected through a 22G needle and swells in situ</td><td>Pain relief from 1 month after surgery for at least 12 months</td><td>CE</td></tr><tr><td>HYADD4-G<sup>85,86</sup> (Fidia Farmaceutici)</td><td>HA 500-730 kDa, functionalized with 2%-3% hexadecylamine</td><td>1.5 mL [8 mg/mL] intradiscal injections guided by X-ray</td><td>Statistically significant pain relief up to 24 weeks</td><td>Clinical Trials</td></tr><tr><td>BioDisc<sup>87,88</sup> (CryoLife)</td><td>Albumin + glutaraldehyde hydrogel</td><td>Crosslink in situ within 2 min</td><td>Reduction in pain after 6 months</td><td>Unknown<sup>c</sup></td></tr><tr><td>NuCore<sup>89,90</sup> (Spine Wave)</td><td>Block polymers of silk and elastin crosslinked in situ with diisocyanate</td><td>Injected with a sealed vented needle to recover disc height (0.3-1.9 mL)</td><td>Reduction in back and leg pain, regained disc height</td><td>Unknown<sup>c</sup></td></tr><tr><td></td><td></td><td></td><td></td><td></td><td></td></tr></tbody></table>

Note: The list of products for viscosupplementation (VS) is non-exhaustive due to many products on the market. Instead, representative products for different materials and crosslinking mechanisms are presented. The other types of products are exhaustive to the authors' best knowledge.

Abbreviations: CaP, calcium phosphate; DBM, demineralized bone matrix; DDD, degenerative disc disease; EMD, enamel matrix derivatives; FDA, Food and Drug Administration; HA, hyaluronic acid; PBS, phosphate-buffered saline; SHA, sodium hyaluronate; VS, viscosupplementation.

<footnote><sup>a</sup>Clinical trial has a 5-year follow-up period.</footnote>

<footnote><sup>b</sup>Bone defects = trauma, oncology, craniofacial.</footnote>

<footnote><sup>c</sup>Seems to be discontinued.</footnote>
effect. In this section, we describe some of the leading clinical products for VS. In addition, we will discuss the products that have undergone clinical trials or been commercialized in the EU or the United States to regenerate bone, cartilage, or NP tissue. The products we will discuss are summarized in Table 1. We present their application indications, therapeutic effect, delivery method, and composition. Apart from VS, the list is exhaustive to the authors' best knowledge but might suffer from lack of data availability as many manufacturers choose to keep data on file rather than publishing their results. With the introduction of the European EUDAMED database, this is expected to change within the EU market. Bone putties (DBM/inorganic particles in hydrogel carrier) have been excluded for bone regeneration products unless they are marketed as injectable gels.

Many manufacturers have chosen to not publish their findings but keep their data privately on file. This applies to the products AphaGRAFT®, Kinex®, AlloFuse®, and Tactoset®, meaning we have limited information on these products which can limit our discussion of these solutions.

## 3.1 | Cartilage treatment

An exciting area where injectable hydrogels have become an established treatment is cartilage degeneration in joints. This is primarily indicated by OA, a disease-causing degeneration of the cartilage and the subchondral bone in the joints and affects roughly a third of people above 65 years,⁹¹ thereby having a high socioeconomic cost. In addition to degenerated cartilage and subchondral bone, synovitis and systemic inflammation are part of the pathogenesis.⁹² Patients with mild to moderate OA usually are treated with intra-

articular injection of corticosteroids, as it provides an anti-inflammatory effect.⁹³ However, corticosteroids are just capable of treating the symptoms, that is, reducing pain, but not able to stop the progress of OA.⁹⁴ Therefore, VS has become a popular treatment alternative as it provides a longer therapeutic effect.⁹⁵

For late-stage OA, arthroplasty is the preferred treatment, where the joint is partially or totally replaced with a prosthesis that is typically made of cobalt chrome or titanium alloys.⁹⁶ An alternative treatment is microfractures to release chondroprogenitor cells to the diseased location, but this tends to form fibrocartilage instead of desired hyaline cartilage.⁹⁷ The fibrocartilage has inferior mechanical properties than the native hyaline cartilage,⁹⁸ providing a temporary solution. Injectable hydrogels have become an attractive strategy for treatment in OA, both for delaying arthroplasty and attempting to regenerate the damaged cartilage toward more native-like cartilage than what can be achieved from microfracture. The two primary therapies, VS and regeneration, are illustrated in Figure 2.

### 3.1.1 | Viscosupplementation

There are multiple solutions based on HA injection into the knee for pain relief through VS (Table 1). There are two generations of VS products. The first generation consisted of HA solutions dissolved in an aqueous solution. The second generation consisted of crosslinked HA. To maintain injectability for the crosslinked gels, some of these are granulated HA gels chunks (typically less than 80 µm) that are mixed in an aqueous solution. This is for instance the case for Anika's Monovisc®, as can be deduced from its patent.⁹⁹ The clinical efficacy of VS therapies is debated. In a more extensive meta-analysis including 89 trials with 12,677 patients involved, they could not observe

<figure><img src="image_4.png" /><figcaption>**FIGURE 2** Treatment of cartilage defects caused by osteoarthritis.

(a) Viscosupplementation using hyaluronic acid to obtain improved joint movement and pain relief. (b) Cartilage regeneration using injectable hydrogels</figcaption></figure>
effect. In this section, we describe some of the leading clinical products for VS. In addition, we will discuss the products that have undergone clinical trials or been commercialized in the EU or the United States to regenerate bone, cartilage, or NP tissue. The products we will discuss are summarized in Table 1. We present their application indications, therapeutic effect, delivery method, and composition. Apart from VS, the list is exhaustive to the authors' best knowledge but might suffer from lack of data availability as many manufacturers choose to keep data on file rather than publishing their results. With the introduction of the European EUDAMED database, this is expected to change within the EU market. Bone putties (DBM/inorganic particles in hydrogel carrier) have been excluded for bone regeneration products unless they are marketed as injectable gels.

Many manufacturers have chosen to not publish their findings but keep their data privately on file. This applies to the products AphaGRAFT®, Kinex®, AlloFuse®, and Tactoset®, meaning we have limited information on these products which can limit our discussion of these solutions.

**3.1 | Cartilage treatment**

An exciting area where injectable hydrogels have become an established treatment is cartilage degeneration in joints. This is primarily indicated by OA, a disease-causing degeneration of the cartilage and the subchondral bone in the joints and affects roughly a third of people above 65 years,⁹¹ thereby having a high socioeconomic cost. In addition to degenerated cartilage and subchondral bone, synovitis and systemic inflammation are part of the pathogenesis.⁹²
Patients with mild to moderate OA usually are treated with intra-

articular injection of corticosteroids, as it provides an anti-inflammatory effect.⁹³ However, corticosteroids are just capable of treating the symptoms, that is, reducing pain, but not able to stop the progress of OA.⁹⁴ Therefore, VS has become a popular treatment alternative as it provides a longer therapeutic effect.⁹⁵

For late-stage OA, arthroplasty is the preferred treatment, where the joint is partially or totally replaced with a prosthesis that is typically made of cobalt chrome or titanium alloys.⁹⁶ An alternative treatment is microfractures to release chondroprogenitor cells to the diseased location, but this tends to form fibrocartilage instead of desired hyaline cartilage.⁹⁷ The fibrocartilage has inferior mechanical properties than the native hyaline cartilage,⁹⁸ providing a temporary solution. Injectable hydrogels have become an attractive strategy for treatment in OA, both for delaying arthroplasty and attempting to regenerate the damaged cartilage toward more native-like cartilage than what can be achieved from microfracture. The two primary therapies, VS and regeneration, are illustrated in Figure 2.

**3.1.1 | Viscosupplementation**

There are multiple solutions based on HA injection into the knee for pain relief through VS (Table 1). There are two generations of VS products. The first generation consisted of HA solutions dissolved in an aqueous solution. The second generation consisted of crosslinked HA. To maintain injectability for the crosslinked gels, some of these are granulated HA gels chunks (typically less than 80 µm) that are mixed in an aqueous solution. This is for instance the case for Anika's Monovisc®, as can be deduced from its patent.⁹⁹ The clinical efficacy of VS therapies is debated. In a more extensive meta-analysis including 89 trials with 12,677 patients involved, they could not observe

<figure><img src="image_4.png" /><figcaption>**FIGURE 2** Treatment of cartilage defects caused by osteoarthritis.
(a) Viscosupplementation using hyaluronic acid to obtain improved joint movement and pain relief. (b) Cartilage regeneration using injectable hydrogels</figcaption></figure>
any clinically relevant benefit.¹⁰⁰ There were, however, indications that high molecular (>6000 kDa) or covalently crosslinked HA could provide a beneficial therapeutic effect.¹⁰⁰ In contrast, in another meta-analysis considering only FDA-approved VS in randomized, saline-controlled trials (29 studies; 4866 patients; active: 2673, control: 2193) they concluded that these products are safe and effective through 26 weeks in patients with symptomatic OA.¹⁰¹ Simultaneously, a consensus of eight European experts on OA discussed the clinical effect of HA in VS: they unanimously agreed that VS is an efficient strategy for managing mild to moderate knee OA, is a cost-efficient treatment in knee OA, but is not an alternative to surgery in advanced hip OA.¹⁰²

Although most of the solutions are based on HA, it is essential to consider the chemical composition and design of the gel. As mentioned before, high molecular weights and covalent crosslinking seem to be preferable. A higher molecular weight HA is believed to have improved residence time and adhesion to the cartilage providing more lasting lubrication under loading,¹⁰³ while a crosslinked HA gel would degrade slower than a non-crosslinked HA solution,¹⁰⁴ giving a longer therapeutic effect. For instance, for a lightly crosslinked VS such as Monovisc®, one injection provides 6 months of therapeutic effect,⁶⁴ compared to three injections for a conventional non-crosslinked VS such as Orthovisc®. This gives significant indirect cost savings in the form of fewer hospitalization visits and reduced pain to the patient. More importantly, it can reduce the occurrence of more serious adverse events such as pseudoseptic reactions (inflammation and swelling of joint without infection, occurs in 1%-3% of patients) that typically occurs after second or third injection.¹⁰⁵

Recent clinical trials demonstrated that injection of HA has anti-inflammatory and antioxidative properties, which can decrease the progression of OA.¹⁰⁶ This effect seems to be mediated through receptor signaling via binding with cluster determinant 44, toll-like receptors 2 and 4, intercellular adhesion molecule I, and layilin, providing a multifactorial mechanism.¹⁰⁷ Additionally, there are indications that high molecular weight HA promotes an anti-inflammatory response, meanwhile, low molecular weight HA favors an inflammatory response.¹⁰⁷ Altogether, intra-articular injections of HA-based VS have demonstrated an effect, and there is still room to tune the hydrogel composition to obtain solutions providing better lubrication with enhanced therapeutic benefit.

A recent commercialization is VS made from polyacrylamide such as Contura's Arthrosemid®. Arthrosemid® is a gel consisting of covalently crosslinked polyacrylamide, which is non-degradable.⁶⁶ It was used initially for veterinary application in horses with OA,¹⁰⁸ but recently the therapeutic effect has been demonstrated to be functional up to 52 weeks in humans.⁶⁷ As the material is non-degradable, the therapeutic effect is expected to be significantly longer. An in vivo subcutaneous rat model comparing the acrylamide gel to a HA gel as soft tissue fillers suggested significantly different in vivo behavior. The acrylamide underwent cell infiltration by macrophages and fibroblasts and tissue integration, meanwhile, cell infiltration did not occur in the HA gel which was encapsulated by a thin fibrous layer.¹⁰⁹

The relevance of the model is limited as the study was conducted in a small animal with a subcutaneous application instead of intra-articular. However, the results may suggest that the clinical mechanisms of HA and acrylamide gels are different.

### 3.1.2 | Cartilage regeneration

Although VS, such as Monovisc® and Orthovisc®, can typically provide pain relief for up to 6 months, they do not regenerate functional cartilage. This has led to an enormous focus on cartilage regeneration, and there is a series of products in clinical trials. They use different tactics for regeneration; conventionally, a microfracture procedure where bone marrow-derived MSCs are released into the defect site has been used for cartilage regeneration, but with considerable variability and inconsistency.¹¹⁰ Both the BST-CarGel® solution and the GelrinC® build on this procedure by providing the released MSCs with a scaffold for guided cartilage regeneration. Their mechanism differs slightly. The BST-CarGel® consists of chitosan dissolved in aqueous glycerol phosphate (buffer at physiological pH), that when mixed with blood forms a clot with increased mechanical properties and longer stability.⁶⁸ The capability of chitosan as a hemostatic agent is derived from its poly(cationic) nature that allows it to bind with the negatively charged thrombocytes and erythrocytes in the blood.¹¹¹ A 5-year follow-up study for treatment of OA in the knee demonstrated significantly better cartilage regeneration with BST-CarGel® compared to microfracture alone,⁶⁹ and their animal trials suggest that the gel also regenerates cartilage with increased hyaline characteristics.¹¹² GelrinC® on the other hand, is based on PEGDA mixed with denatured human fibrinogen and can be injected in liquid form but solidifies into a gel upon 90 s of UVA irradiation.⁷² It can be used for both chondral and osteochondral lesions and showed statistical improvement compared to the absence of treatment after 24-months follow-up.⁷¹ Their MRI data suggested a zonal variation in the cartilage, which they interpret as the cartilage might be hyaline-like rather than fibrous.

Although some indications, neither of the solutions has proven to produce native-like hyaline cartilage in humans. Part of the reason they cannot prove it is that one cannot take histology samples from living patients. Instead, they must use methods such as magnetic resonance imaging (MRI). Unfortunately, clinical MRIs tend to have a moderate resolution, limiting some of the quality of the data used in the analysis.

Although GelinC® and BST-CarGel® have shown short-term improvement, the success is governed by the long-term results, economic viability, and clear improvement from microfracture alone. Frappier and colleagues¹¹³ demonstrated this by evaluating the economic value of BST-CarGel® solution versus microfracture alone using Germany as a reference market. Their results suggest that a positive investment return is reached after 4 years and more than €6400 of cost saved over a 20-year period. Some essential limitations to this study are a lack of long-term clinical data for BST-CarGel® versus microfracture, and it only considers cost and not the quality of life of
any clinically relevant benefit.¹⁰⁰ There were, however, indications that high molecular (>6000 kDa) or covalently crosslinked HA could provide a beneficial therapeutic effect.¹⁰⁰ In contrast, in another meta-analysis considering only FDA-approved VS in randomized, saline-controlled trials (29 studies; 4866 patients; active: 2673, control: 2193) they concluded that these products are safe and effective through 26 weeks in patients with symptomatic OA.¹⁰¹ Simultaneously, a consensus of eight European experts on OA discussed the clinical effect of HA in VS: they unanimously agreed that VS is an efficient strategy for managing mild to moderate knee OA, is a cost-efficient treatment in knee OA, but is not an alternative to surgery in advanced hip OA.¹⁰²

Although most of the solutions are based on HA, it is essential to consider the chemical composition and design of the gel. As mentioned before, high molecular weights and covalent crosslinking seem to be preferable. A higher molecular weight HA is believed to have improved residence time and adhesion to the cartilage providing more lasting lubrication under loading,¹⁰³ while a crosslinked HA gel would degrade slower than a non-crosslinked HA solution,¹⁰⁴ giving a longer therapeutic effect. For instance, for a lightly crosslinked VS such as Monovisc®, one injection provides 6 months of therapeutic effect,⁶⁴ compared to three injections for a conventional non-crosslinked VS such as Orthovisc®. This gives significant indirect cost savings in the form of fewer hospitalization visits and reduced pain to the patient. More importantly, it can reduce the occurrence of more serious adverse events such as pseudoseptic reactions (inflammation and swelling of joint without infection, occurs in 1%-3% of patients) that typically occurs after second or third injection.¹⁰⁵

Recent clinical trials demonstrated that injection of HA has anti-inflammatory and antioxidative properties, which can decrease the progression of OA.¹⁰⁶ This effect seems to be mediated through receptor signaling via binding with cluster determinant 44, toll-like receptors 2 and 4, intercellular adhesion molecule I, and layilin, providing a multifactorial mechanism.¹⁰⁷ Additionally, there are indications that high molecular weight HA promotes an anti-inflammatory response, meanwhile, low molecular weight HA favors an inflammatory response.¹⁰⁷ Altogether, intra-articular injections of HA-based VS have demonstrated an effect, and there is still room to tune the hydrogel composition to obtain solutions providing better lubrication with enhanced therapeutic benefit.

A recent commercialization is VS made from polyacrylamide such as Contura's Arthrosemid®. Arthrosemid® is a gel consisting of covalently crosslinked polyacrylamide, which is non-degradable.⁶⁶ It was used initially for veterinary application in horses with OA,¹⁰⁸ but recently the therapeutic effect has been demonstrated to be functional up to 52 weeks in humans.⁶⁷ As the material is non-degradable, the therapeutic effect is expected to be significantly longer. An in vivo subcutaneous rat model comparing the acrylamide gel to a HA gel as soft tissue fillers suggested significantly different in vivo behavior. The acrylamide underwent cell infiltration by macrophages and fibroblasts and tissue integration, meanwhile, cell infiltration did not occur in the HA gel which was encapsulated by a thin fibrous layer.¹⁰⁹

The relevance of the model is limited as the study was conducted in a small animal with a subcutaneous application instead of intra-articular. However, the results may suggest that the clinical mechanisms of HA and acrylamide gels are different.

### 3.1.2 | Cartilage regeneration

Although VS, such as Monovisc® and Orthovisc®, can typically provide pain relief for up to 6 months, they do not regenerate functional cartilage. This has led to an enormous focus on cartilage regeneration, and there is a series of products in clinical trials. They use different tactics for regeneration; conventionally, a microfracture procedure where bone marrow-derived MSCs are released into the defect site has been used for cartilage regeneration, but with considerable variability and inconsistency.¹¹⁰ Both the BST-CarGel® solution and the GelrinC® build on this procedure by providing the released MSCs with a scaffold for guided cartilage regeneration. Their mechanism differs slightly. The BST-CarGel® consists of chitosan dissolved in aqueous glycerol phosphate (buffer at physiological pH), that when mixed with blood forms a clot with increased mechanical properties and longer stability.⁶⁸ The capability of chitosan as a hemostatic agent is derived from its poly(cationic) nature that allows it to bind with the negatively charged thrombocytes and erythrocytes in the blood.¹¹¹ A 5-year follow-up study for treatment of OA in the knee demonstrated significantly better cartilage regeneration with BST-CarGel® compared to microfracture alone,⁶⁹ and their animal trials suggest that the gel also regenerates cartilage with increased hyaline characteristics.¹¹² GelrinC® on the other hand, is based on PEGDA mixed with denatured human fibrinogen and can be injected in liquid form but solidifies into a gel upon 90 s of UVA irradiation.⁷² It can be used for both chondral and osteochondral lesions and showed statistical improvement compared to the absence of treatment after 24-months follow-up.⁷¹ Their MRI data suggested a zonal variation in the cartilage, which they interpret as the cartilage might be hyaline-like rather than fibrous.

Although some indications, neither of the solutions has proven to produce native-like hyaline cartilage in humans. Part of the reason they cannot prove it is that one cannot take histology samples from living patients. Instead, they must use methods such as magnetic resonance imaging (MRI). Unfortunately, clinical MRIs tend to have a moderate resolution, limiting some of the quality of the data used in the analysis.

Although GelinC® and BST-CarGel® have shown short-term improvement, the success is governed by the long-term results, economic viability, and clear improvement from microfracture alone. Frappier and colleagues¹¹³ demonstrated this by evaluating the economic value of BST-CarGel® solution versus microfracture alone using Germany as a reference market. Their results suggest that a positive investment return is reached after 4 years and more than €6400 of cost saved over a 20-year period. Some essential limitations to this study are a lack of long-term clinical data for BST-CarGel® versus microfracture, and it only considers cost and not the quality of life of
the patient. Nevertheless, the data suggest that it is clinically feasible
to use these different solutions along with microfracture. This should
motivate other research to develop new solutions with improved effi-
cacy and at lower costs. A key challenge the field should address is
successfully regenerating native-like hyaline cartilage and developing
non-invasive methods that can aid in its characterization in vivo. Most
likely some type of agents, such as microfibers or a biomolecule, is
required to guide the direction of the tissue regrowth. Furthermore,
regrowth should preferably follow the zonal tissue architecture that
can be observed in the native articulate cartilage.¹¹⁴,¹¹⁵ Ideally, the
cartilage should recruit chondrocytes or MSCs without the need for
autologous chondrocyte transplantation or microfracture, but there
are currently no such solutions to the authors' best knowledge. At the
time of this review, microfracture procedures are estimated to cost
€4329 and autologous chondrocyte implantation €14,238.¹¹⁶ On top
of this comes the cost of the hydrogel used. Hydrogel scaffolds that
can induce regeneration using only locally recruited chondrocytes can
provide considerable cost savings through reduced surgery times and
trauma to the patients.

Another trend that starts to arise is VS-like products with addi-
tional regenerative capabilities. An example of this is ProCore's Reg-
enoGel® solution that was commercially approved in Israel in 2016
and recently completed their FDA phase 4 clinical trials. RegenoGel®
is based on HA that is mixed with purified platelet-rich plasma-derived
fibrinogen that conjugates to form an injectable gel.⁷⁰ In their clinical
trials, they have been able to demonstrate that the gel is efficient at
treating the symptoms of OA, that is, pain and knee stiffness, for at
least 6 months after treatment start,⁷⁰ but more detailed studies are
required to investigate the long-term effect and the ability to regener-
ate cartilage. Nevertheless, their in vivo cartilage-bone explant mouse
model suggests that the material recruits endogenous cells and differ-
entiates them toward a chondrocyte lineage, yielding significant depo-
sition of GAG-proteins and collagen type 2.¹¹⁷ Although promising for
cartilage regeneration, they have yet to demonstrate cartilage regen-
eration in humans.

## 3.2 | Bone regeneration

Healthy bone is vital for structural stability in the musculoskeletal sys-
tem, and defects result in pain, disability, and reduced mobility in indi-
viduals. Additionally, the treatment of bone defects is a tremendous
burden to healthcare providers, estimating an annual cost of $5 billion
in the United States alone.¹¹⁸ Even though bone defects are rarely
directly mortal, the trauma-induced can be hard to recover from. If we
consider the case of hip fractures, for elder women (>65 years) there
is a 10% likelihood of mortality within 3 months of a hip fracture.¹¹⁹
Similarly, a larger meta-analysis demonstrated that the risk of mortal-
ity is increased by a 6- and 8-fold the first 3 months after hip fracture
for older women and men (>50 years), respectively.¹²⁰ Nor are there
any good treatment alternatives in these cases. In fact, another meta-
analysis demonstrated that the mortality rate 1 year after hip fracture
surgery is 24.5%,¹²¹ suggesting two scenarios: (1) the current medical

devices do not have an appropriate therapeutic effect for the elderly
population, or (2) the current surgical procedure's invasiveness leads
to a challenging recovery for elderly patients.

Bone defects can be widely different, and the products used
depend on defect size and loading level.¹²² Therefore, this section has
been split into three subsections: (1) dental and maxillofacial,
(2) trauma and oncology, and (3) spinal fusion. We treat spinal fusion
as a separate application as it is the largest application area of bone
grafts measured according to market value¹²³ and compared with the
dental and traumatic and oncologic applications, this is a form of het-
erotopic ossification. Hydrogels for bone regeneration has been illus-
trated for two indications in Figure 3.

### 3.2.1 | Dental and maxillofacial

Straumann's Emdogain® dominates the dental market and has more
than 20 years of clinical documentation.¹²⁴ Emdogain® is based on a
porcine enamel matrix derivative, a cocktail of proteins consisting of
amelogenin (90%) and a few other nanomelogenin such as
ameloblastin, enamelin, and tuftelin, carried in an aqueous gel solution
composed of propylene glycol alginate.³⁹ Several of these proteins are
identified as intrinsically disordered polypeptides with a one-to-many
signaling effects in vivo and allow for the formation of multiple tissues
in the injection location.¹²⁵ Emdogain® has been proven to regenerate
multiple periodontal tissues, including the osseo-like tissues, acellular
cementum,¹²⁶ and alveolar bone,¹²⁷ in addition to connective tissues
such as periodontal ligament.¹²⁸ The details of the therapeutic effect
of Emdogain® have been discussed in detail in our former review.³⁹ A
limitation worth noting with Emdogain® is that since it is physically
crosslinked, the degradation occurs quicker than it would with a cova-
lently crosslinked hydrogel. The consequence of this is that the
mechanical properties degrade quickly, and it can no longer keep the
soft tissue flap up, causing a collapse of the gel and limiting the space
available for bone regeneration.¹²⁹

Another product that is well established in the dental domain is
NovaBone's Perioglas® putty. Initially, it was commercialized as a
moldable putty, but a syringe and a cartridge injection system have
since been developed. The gel-like putty consists of calcium sodium
phosphosilicate, more specifically Bioglass® 45S5 particles of 32-
710 µm diameter, delivered through a gel-like binder of PEG.⁷⁶ The
binder is water-soluble and is resorbed within 48-72 h after
implantation,⁷⁶ hence it is the Bioglass® that has the main therapeutic
effect. According to Jones,¹³⁰ the Bioglass® draws its bioactivity from
two mechanisms: (1) the accumulation of glass dissolution products
provides nucleation sites for a hydroxycarbonate apatite layer that
bonds to the surrounding bone. This layer also allows the protein to
attach and cells to attach, proliferate and produce ECM; (2) the
release of dissolution products also plays an active role in driving
osteogenesis through guiding osteoprogenitor cells down an osteo-
blastic differentiation path, and the osteoblasts are transitioned
from a resting stage (G0) to a growth stage (G1). There are, how-
ever, concerns regarding inflammatory foreign body reaction
the patient. Nevertheless, the data suggest that it is clinically feasible
to use these different solutions along with microfracture. This should
motivate other research to develop new solutions with improved effi-
cacy and at lower costs. A key challenge the field should address is
successfully regenerating native-like hyaline cartilage and developing
non-invasive methods that can aid in its characterization in vivo. Most
likely some type of agents, such as microfibers or a biomolecule, is
required to guide the direction of the tissue regrowth. Furthermore,
regrowth should preferably follow the zonal tissue architecture that
can be observed in the native articulate cartilage.¹¹⁴,¹¹⁵ Ideally, the
cartilage should recruit chondrocytes or MSCs without the need for
autologous chondrocyte transplantation or microfracture, but there
are currently no such solutions to the authors' best knowledge. At the
time of this review, microfracture procedures are estimated to cost
€4329 and autologous chondrocyte implantation €14,238.¹¹⁶ On top
of this comes the cost of the hydrogel used. Hydrogel scaffolds that
can induce regeneration using only locally recruited chondrocytes can
provide considerable cost savings through reduced surgery times and
trauma to the patients.

Another trend that starts to arise is VS-like products with addi-
tional regenerative capabilities. An example of this is ProCore's Reg-
enoGel® solution that was commercially approved in Israel in 2016
and recently completed their FDA phase 4 clinical trials. RegenoGel®
is based on HA that is mixed with purified platelet-rich plasma-derived
fibrinogen that conjugates to form an injectable gel.⁷⁰ In their clinical
trials, they have been able to demonstrate that the gel is efficient at
treating the symptoms of OA, that is, pain and knee stiffness, for at
least 6 months after treatment start,⁷⁰ but more detailed studies are
required to investigate the long-term effect and the ability to regener-
ate cartilage. Nevertheless, their in vivo cartilage-bone explant mouse
model suggests that the material recruits endogenous cells and differ-
entiates them toward a chondrocyte lineage, yielding significant depo-
sition of GAG-proteins and collagen type 2.¹¹⁷ Although promising for
cartilage regeneration, they have yet to demonstrate cartilage regen-
eration in humans.

## 3.2 | Bone regeneration

Healthy bone is vital for structural stability in the musculoskeletal sys-
tem, and defects result in pain, disability, and reduced mobility in indi-
viduals. Additionally, the treatment of bone defects is a tremendous
burden to healthcare providers, estimating an annual cost of $5 billion
in the United States alone.¹¹⁸ Even though bone defects are rarely
directly mortal, the trauma-induced can be hard to recover from. If we
consider the case of hip fractures, for elder women (>65 years) there
is a 10% likelihood of mortality within 3 months of a hip fracture.¹¹⁹
Similarly, a larger meta-analysis demonstrated that the risk of mortal-
ity is increased by a 6- and 8-fold the first 3 months after hip fracture
for older women and men (>50 years), respectively.¹²⁰ Nor are there
any good treatment alternatives in these cases. In fact, another meta-
analysis demonstrated that the mortality rate 1 year after hip fracture
surgery is 24.5%,¹²¹ suggesting two scenarios: (1) the current medical

devices do not have an appropriate therapeutic effect for the elderly
population, or (2) the current surgical procedure's invasiveness leads
to a challenging recovery for elderly patients.

Bone defects can be widely different, and the products used
depend on defect size and loading level.¹²² Therefore, this section has
been split into three subsections: (1) dental and maxillofacial,
(2) trauma and oncology, and (3) spinal fusion. We treat spinal fusion
as a separate application as it is the largest application area of bone
grafts measured according to market value¹²³ and compared with the
dental and traumatic and oncologic applications, this is a form of het-
erotopic ossification. Hydrogels for bone regeneration has been illus-
trated for two indications in Figure 3.

### 3.2.1 | Dental and maxillofacial

Straumann's Emdogain® dominates the dental market and has more
than 20 years of clinical documentation.¹²⁴ Emdogain® is based on a
porcine enamel matrix derivative, a cocktail of proteins consisting of
amelogenin (90%) and a few other nanomelogenin such as
ameloblastin, enamelin, and tuftelin, carried in an aqueous gel solution
composed of propylene glycol alginate.³⁹ Several of these proteins are
identified as intrinsically disordered polypeptides with a one-to-many
signaling effects in vivo and allow for the formation of multiple tissues
in the injection location.¹²⁵ Emdogain® has been proven to regenerate
multiple periodontal tissues, including the osseo-like tissues, acellular
cementum,¹²⁶ and alveolar bone,¹²⁷ in addition to connective tissues
such as periodontal ligament.¹²⁸ The details of the therapeutic effect
of Emdogain® have been discussed in detail in our former review.³⁹ A
limitation worth noting with Emdogain® is that since it is physically
crosslinked, the degradation occurs quicker than it would with a cova-
lently crosslinked hydrogel. The consequence of this is that the
mechanical properties degrade quickly, and it can no longer keep the
soft tissue flap up, causing a collapse of the gel and limiting the space
available for bone regeneration.¹²⁹

Another product that is well established in the dental domain is
NovaBone's Perioglas® putty. Initially, it was commercialized as a
moldable putty, but a syringe and a cartridge injection system have
since been developed. The gel-like putty consists of calcium sodium
phosphosilicate, more specifically Bioglass® 45S5 particles of 32-
710 µm diameter, delivered through a gel-like binder of PEG.⁷⁶ The
binder is water-soluble and is resorbed within 48-72 h after
implantation,⁷⁶ hence it is the Bioglass® that has the main therapeutic
effect. According to Jones,¹³⁰ the Bioglass® draws its bioactivity from
two mechanisms: (1) the accumulation of glass dissolution products
provides nucleation sites for a hydroxycarbonate apatite layer that
bonds to the surrounding bone. This layer also allows the protein to
attach and cells to attach, proliferate and produce ECM; (2) the
release of dissolution products also plays an active role in driving
osteogenesis through guiding osteoprogenitor cells down an osteo-
blastic differentiation path, and the osteoblasts are transitioned
from a resting stage (G0) to a growth stage (G1). There are, how-
ever, concerns regarding inflammatory foreign body reaction
<figure><img src="image_5.png" /><figcaption>**FIGURE 3** Illustration of hydrogel application for bone regeneration. Top panel: fracture healing in traumatology using a needle-injected hydrogel. Bottom panel: spinal fusion using arthroscopic injection of a ceramic particle loaded hydrogel</figcaption></figure>

around the bioglass particles that might limit the clinical success of the putty.¹³¹

### 3.2.2 | Orthopedics: Trauma and oncology

An approved product for orthopedics is the Baxter Actifuse® Flow. It consists of silicon substituted calcium phosphate particles of size 90–500 µm carried in an aqueous polymer carrier consisting of poloxamer 407 (P407).77,78 The P407 is a triblock polymer with a hydrophobic polypropylene glycol core and hydrophilic PEG side arms, that goes through a thermoreversible gelation mechanism, meaning that the solution gels above a given temperature.¹³² The temperature for which the sol–gel transition occurs decreases with the P407 concentration, and it has been demonstrated that for a concentration of 16.5% (wt.% in purified water) the solution gels at a temperature of 27.1°C.¹³³ It can be speculated that the Actifuse® Flow carrier has a P407 concentration of 16.5 wt.% or lower, meaning that it will be liquid at room temperature while at physiological temperatures it would form a gel. The solution has successfully treated benign bone defects in the pediatric population.¹³⁴ A series of similar solutions has been made combining demineralized bone matrix (DBM) particles in similar

reverse-phase medium-based hydrogels. This includes both Dynagraft® III and AlphaGRAFT® that combines DBM particles with poloxamer gel, meanwhile, AlloFuse® combines DBM particles in a carrier of polyethylene oxide polypropylene oxide block copolymer. Optium® and Grafton® DBM uses glycol as a carrier. Unfortunately, with the exception of Medtronic with their Grafton® product, these manufacturers have chosen to keep their data on file so the products cannot be discussed directly. In general terms, DBM is an attractive biomaterial as the acid-extraction process allows the retention of growth factors such as bone morphogenic proteins (BMPs), yielding osteoinductive properties, but it is a challenge for manufacturers to sterilize DBMs without inactivating these growth factors.¹³⁵ Due to the risk of immunoreactions and transmission of infections, the use of DBM and other allograft products is regulatorily unfavorable in Europe, and with the new MDR, it is expected to be limited further.¹³⁶

A more recent solution is Anika's Tactoset® solution where calcium phosphate particles and HA are mixed into a hardening, injectable gel solution.⁸² Currently, it has only been published as a technical note with limited information on the composition and therapeutic effect. A similar solution is Globus Medical's Kinex® composed of bioglass and collagen in a HA gel.⁸³ However, the manufacturer has
<figure><img src="image_5.png" /><figcaption>**FIGURE 3** Illustration of hydrogel application for bone regeneration. Top panel: fracture healing in traumatology using a needle-injected hydrogel. Bottom panel: spinal fusion using arthroscopic injection of a ceramic particle loaded hydrogel</figcaption></figure>

around the bioglass particles that might limit the clinical success of
the putty.¹³¹

3.2.2 | Orthopedics: Trauma and oncology

An approved product for orthopedics is the Baxter Actifuse® Flow. It consists of silicon substituted calcium phosphate particles of size 90–500 µm carried in an aqueous polymer carrier consisting of poloxamer 407 (P407).<sup>77,78</sup> The P407 is a triblock polymer with a hydrophobic polypropylene glycol core and hydrophilic PEG side arms, that goes through a thermoreversible gelation mechanism, meaning that the solution gels above a given temperature.<sup>132</sup> The temperature for which the sol–gel transition occurs decreases with the P407 con- centration, and it has been demonstrated that for a concentration of 16.5% (wt.% in purified water) the solution gels at a temperature of 27.1°C.<sup>133</sup> It can be speculated that the Actifuse® Flow carrier has a P407 concentration of 16.5 wt.% or lower, meaning that it will be liq- uid at room temperature while at physiological temperatures it would form a gel. The solution has successfully treated benign bone defects in the pediatric population.<sup>134</sup> A series of similar solutions has been made combining demineralized bone matrix (DBM) particles in similar

reverse-phase medium-based hydrogels. This includes both
Dynagraft® III and AlphaGRAFT® that combines DBM particles with
poloxamer gel, meanwhile, AlloFuse® combines DBM particles in a
carrier of polyethylene oxide polypropylene oxide block copolymer.
Optium® and Grafton® DBM uses glycol as a carrier. Unfortunately,
with the exception of Medtronic with their Grafton® product, these
manufacturers have chosen to keep their data on file so the products
cannot be discussed directly. In general terms, DBM is an attractive
biomaterial as the acid-extraction process allows the retention of
growth factors such as bone morphogenic proteins (BMPs), yielding
osteoinductive properties, but it is a challenge for manufacturers to
sterilize DBMs without inactivating these growth factors.¹³⁵ Due to
the risk of immunoreactions and transmission of infections, the use of
DBM and other allograft products is regulatorily unfavorable in
Europe, and with the new MDR, it is expected to be limited
further.¹³⁶

A more recent solution is Anika's Tactoset® solution where calci-
cium phosphate particles and HA are mixed into a hardening, inject-
able gel solution.⁸² Currently, it has only been published as a technical
note with limited information on the composition and therapeutic
effect. A similar solution is Globus Medical's Kinex® composed of bio-
glass and collagen in a HA gel.⁸³ However, the manufacturer has
chosen to keep their data on file; hence no research is published on
this solution.

3.2.3 | Orthopedics: Spinal fusion

Spinal fusion is a common surgery requiring bone-growing implants,
with approximately 200,000 lumbar spinal fusions conducted in 2015
in the United States alone.¹³⁷ Spinal fusion is performed to compen-
sate for degenerative disc disease (DDD) where the height of inter-
vertebral disc (IVD) has reduced leading to the compression of the
spinal cord, which translates to back pain. Degenerative disc, the first
step toward DDD, affects more than 90% of people above
50 years.¹³⁸ When the degeneration progresses, spinal fusion is an
attractive surgery for pain mitigation and preventing damage to the
spinal cord. The surgery typically consists of a cage being inserted to
mechanically regain the spacing between the vertebrae, then bone
grafts are used to stimulate bone growth to fuse together the adjacent
vertebras. Conventionally an open surgery is used, but there is now a
trend to use minimally invasive procedures (MIP) such as key-hole sur-
gery.¹³⁹ MIP can be incompatible with conventional bone grafts due
to large size or high viscosity; hence this trend favors injectable solu-
tions such as hydrogels. Moreover, MIP spinal fusion requires less
bone to be removed for access to the IVD, which means less

autologous bone available as graft material, increasing the demand for
alternative grafting materials. Between 9% and 39% of lumbar spinal
fusions fail,¹⁴⁰ indicating a need for more potent bone regrowth solu-
tions. Spinal fusion requires heterotopic ossification, meaning bone
tissue growth in soft-tissue locations where bone is usually not pre-
sent. This makes it a challenging task, and a graft only exhibiting
osteoconductive properties is suboptimal. Ideally, for treatment of
large defects and for heterotopic ossification the graft should be oste-
oinductive, a phenomenon induced when the material creates a local
homeostatic imbalance by binding to calcium and/or phosphate ions,
causing depletion of these ions.¹⁴¹,¹⁴² Hence, an osteoinductive mate-
rial is likely to quickly induce a stable fusion than a graft that is just
osteoconductive. This has motivated many to introduce BMP in their
graft products, for example, Medtronic uses rhBMP-2 in their Infuse®
(US)/Induct® (EU) bone graft. However, according to the European
regulation, the BMP makes it considered a medicinal product. Further-
more, the use of rhBMP-2 in this product has been linked to several
adverse events where the high doses of the growth factor, mainly
when used for “off-label” cervical spinal surgeries, causes an inflam-
matory effect yielding high complication rates.¹⁴³ A similar BMP-7
based product named OP-1® from Stryker has failed to obtain FDA
approval for similar spinal applications. However, effective hydrogel
therapies are emerging that do not depend on BMP-based growth
factors to obtain their therapeutic effect. The before mentioned

<figure><img src="image_6.png" /><figcaption>FIGURE 4 (a) Illustration of the intervertebral disc. (b) Strain and stress levels in non-degenerated and degenerated intervertebral discs demonstrating how the degenerated disc is prone to higher stress levels, particularly around the AF region. Reproduced and adapted from 137, 138. Reprinted with permission from copyright CC BY 4.0. AF, annulus fibrosus; BEP, bony endplate; CEP, cartilaginous endplate; NP, nucleus pulposus</figcaption></figure>
chosen to keep their data on file; hence no research is published on
this solution.

3.2.3 | Orthopedics: Spinal fusion

Spinal fusion is a common surgery requiring bone-growing implants,
with approximately 200,000 lumbar spinal fusions conducted in 2015
in the United States alone.¹³⁷ Spinal fusion is performed to compen-
sate for degenerative disc disease (DDD) where the height of inter-
vertebral disc (IVD) has reduced leading to the compression of the
spinal cord, which translates to back pain. Degenerative disc, the first
step toward DDD, affects more than 90% of people above
50 years.¹³⁸ When the degeneration progresses, spinal fusion is an
attractive surgery for pain mitigation and preventing damage to the
spinal cord. The surgery typically consists of a cage being inserted to
mechanically regain the spacing between the vertebrae, then bone
grafts are used to stimulate bone growth to fuse together the adjacent
vertebras. Conventionally an open surgery is used, but there is now a
trend to use minimally invasive procedures (MIP) such as key-hole sur-
gery.¹³⁹ MIP can be incompatible with conventional bone grafts due
to large size or high viscosity; hence this trend favors injectable solu-
tions such as hydrogels. Moreover, MIP spinal fusion requires less
bone to be removed for access to the IVD, which means less

autologous bone available as graft material, increasing the demand for
alternative grafting materials. Between 9% and 39% of lumbar spinal
fusions fail,¹⁴⁰ indicating a need for more potent bone regrowth solu-
tions. Spinal fusion requires heterotopic ossification, meaning bone
tissue growth in soft-tissue locations where bone is usually not pre-
sent. This makes it a challenging task, and a graft only exhibiting
osteoconductive properties is suboptimal. Ideally, for treatment of
large defects and for heterotopic ossification the graft should be oste-
oinductive, a phenomenon induced when the material creates a local
homeostatic imbalance by binding to calcium and/or phosphate ions,
causing depletion of these ions.¹⁴¹,¹⁴² Hence, an osteoinductive mate-
rial is likely to quickly induce a stable fusion than a graft that is just
osteoconductive. This has motivated many to introduce BMP in their
graft products, for example, Medtronic uses rhBMP-2 in their Infuse®
(US)/Induct® (EU) bone graft. However, according to the European
regulation, the BMP makes it considered a medicinal product. Further-
more, the use of rhBMP-2 in this product has been linked to several
adverse events where the high doses of the growth factor, mainly
when used for “off-label” cervical spinal surgeries, causes an inflam-
matory effect yielding high complication rates.¹⁴³ A similar BMP-7
based product named OP-1® from Stryker has failed to obtain FDA
approval for similar spinal applications. However, effective hydrogel
therapies are emerging that do not depend on BMP-based growth
factors to obtain their therapeutic effect. The before mentioned

<figure><img src="image_6.png" /><figcaption>FIGURE 4 (a) Illustration of the intervertebral disc. (b) Strain and stress levels in non-degenerated and degenerated intervertebral discs demonstrating how the degenerated disc is prone to higher stress levels, particularly around the AF region. Reproduced and adapted from 137, 138. Reprinted with permission from copyright CC BY 4.0. AF, annulus fibrosus; BEP, bony endplate; CEP, cartilaginous endplate; NP, nucleus pulposus</figcaption></figure>
Baxter Actifuse® Flow has successfully been used for spinal fusion procedure.¹⁴⁴ When used in a comparative clinical study to the Medtronic Infuse® graft, they were able to demonstrate similar fusion rates (Actifuse® 9/9, Infuse® 8/9 cases) and both products yielded similar alleviation of pain and improved quality of life.¹⁴⁵ Also, DBM solutions have been approved clinically for spinal application. The Grafton® DBM was tested in a clinical trial with a total of 120 patients undergoing posterolateral spinal fusion, of which 81 (70%) completed the 24-month radiographic study.⁸⁰ Grafton® was used on one side of the spine and autograft on the other, and in 42 (52%) of the Grafton® cases, successful fusion was obtained versus 44 (54%) for the autograft side. The authors concluded that the Grafton® DBM gel can be used to extend autograft material during spinal fusion.

## 3.3 | Nucleus pulposus

Spinal fusion tends to be conducted due to DDD, where the IVD has degraded and lost its height or fractured. The IVD is to find between all the vertebra of the spine. It has three main components; the hydrogel-like NP in the core, surrounded by the annulus fibrosus (AF), and cartilaginous end plates (CEP) at the top and bottom (Figure 4a).¹⁴⁶

The NP consists of approximately 50% (dry weight) proteoglycan proteins that play a vital role in binding water in the NP and shock absorbance.¹⁴⁷ During disc degradation, the concentration of proteoglycans decreases, causing a drop in stiffness.¹⁴⁶ This increases the risk of AF bulging, increases the compressive strain on the AF (Figure 4b), and increases the chances of peripheral failure of the end plates.¹⁴⁸ Therefore, a potential treatment of DDD would be to repair the NP.

A solution that has been approved for the European market is GelStix®. GelStix® uses a dehydrated polyacrylonitrile that is injected into the NP through a 22-G needle in the form of a filament, where it gets hydrated from the surrounding body liquids and expands 10-fold.⁸⁴ In a 12-month follow-up with 29 patients, 86.2% rated the procedure as very good or good, and pain relief was observed already after 1 month.⁸⁴ However, there have been reported complications associated with this procedure. Durdag and colleagues reoperated a woman with a GelStix® implanted as she was admitted with severe radicular pain.¹⁴⁹ The pain was linked to a fragment of implant that had penetrated through an annual tear and caused compression to the spinal root. The authors speculate that the implant may have been initially wrongly placed in the AF, highlighting the importance of the correct placement of the implant.

HA with a similar composition to the solutions used for VS has been used for treatment of the NP. In a 24-week follow-up period, Mazza and co-workers observed relief from chronic lower back pain due to DDD compared to the baseline.⁸⁵ They had two patients drop out due to adverse events, but this is not believed to be related to the treatment. However, the clinical efficacy is proven only over a short time period. Considering the surgical risk related to bypassing vital organs during injection, this therapy can come short when evaluating

it using a cost-benefit analysis. Hence a longer-lasting therapy should be investigated.

Two other solutions have been tried clinically, but seem to have been discontinued. The NuCore® gel for NP replacement consists of elastin and silk co-polymers that are crosslinked in situ.⁸⁹ A 2-year follow-up pilot clinical study with 14 patients demonstrated a significant reduction in back and leg pain, regained disc height, and no side effects.⁸⁹ There have not been any clinical publications on this product since 2009, and it seems to have been discontinued by the supplier. CryoLife started clinical trials on their product BioDisc but have not published the outcome of the trial. In a conference abstract containing interim results, they reported at the 6-month follow-up a decrease in mean Oswestry Disability index from 49.2 to 11 and a decrease in numerical pain score from 5.86 to 1.62,⁸⁸ which could seem promising. However, they also reported that 2 of the 10 patients enrolled experienced recurrent herniation requiring surgery. After this abstract from 2008, there has been no publication, and the product seems to have been discontinued.

Since neither CryoLife nor Spine Wave has disclosed why their products were discontinued, it is not feasible to conclude why they failed to perform in the clinic.

## 4 | REGULATORY CLASSIFICATION AND CONSEQUENCES

From a regulatory perspective, the first step of translating a medical device is to assign it to the appropriate risk classification group, namely risk class. In Europe with the new MDR, this is reasonably straightforward with injectable hydrogels. Because they are implanted, hence in contact with human tissue over a prolonged period and have a biological effect, it becomes a class III device (highest risk level). This means a premarket clinical investigation is mandatory. This can be mitigated if equivalence to a predicate device can be demonstrated. Nevertheless, appropriate equivalence is practically infeasible unless the manufacturer of the new device either (a) also manufactures the predicate device or (b) has a contractual agreement with the manufacturer of the predicate to access all technical information. In the United States, the risk classification differs from Europe as it depends on product device groups. VS products or dental biologics (e.g., Emdogain®) are class III (highest risk), meanwhile more conventional bone graft materials without human growth factors such as Anika's Tactoset® or the DBM solutions are class II. For class II and some class III product groups the 510(k)-pathway can be used if it demonstrates substantial equivalence with existing approved devices, demonstrating that the device is safe and efficient, which is significantly cheaper than introducing a new device. In the case of class III, the 510(k) allows the manufacturer to partially bypass the premarket approval application, meaning they do not need to run a clinical investigation, but this is not applicable for the VS products or dental biological materials discussed here. When the 510(k) is not applicable for the class III devices, the product needs to be evaluated on a case-by-case basis by the authorities (US-FDA).
Baxter Actifuse® Flow has successfully been used for spinal fusion procedure.¹⁴⁴ When used in a comparative clinical study to the Medtronic Infuse® graft, they were able to demonstrate similar fusion rates (Actifuse® 9/9, Infuse® 8/9 cases) and both products yielded similar alleviation of pain and improved quality of life.¹⁴⁵ Also, DBM solutions have been approved clinically for spinal application. The Grafton® DBM was tested in a clinical trial with a total of 120 patients undergoing posterolateral spinal fusion, of which 81 (70%) completed the 24-month radiographic study.⁸⁰ Grafton® was used on one side of the spine and autograft on the other, and in 42 (52%) of the Grafton® cases, successful fusion was obtained versus 44 (54%) for the autograft side. The authors concluded that the Grafton® DBM gel can be used to extend autograft material during spinal fusion.

## 3.3 | Nucleus pulposus

Spinal fusion tends to be conducted due to DDD, where the IVD has degraded and lost its height or fractured. The IVD is to find between all the vertebra of the spine. It has three main components; the hydrogel-like NP in the core, surrounded by the annulus fibrosus (AF), and cartilaginous end plates (CEP) at the top and bottom (Figure 4a).¹⁴⁶

The NP consists of approximately 50% (dry weight) proteoglycan proteins that play a vital role in binding water in the NP and shock absorbance.¹⁴⁷ During disc degradation, the concentration of proteoglycans decreases, causing a drop in stiffness.¹⁴⁶ This increases the risk of AF bulging, increases the compressive strain on the AF (Figure 4b), and increases the chances of peripheral failure of the end plates.¹⁴⁸ Therefore, a potential treatment of DDD would be to repair the NP.

A solution that has been approved for the European market is GelStix®. GelStix® uses a dehydrated polyacrylonitrile that is injected into the NP through a 22-G needle in the form of a filament, where it gets hydrated from the surrounding body liquids and expands 10-fold.⁸⁴ In a 12-month follow-up with 29 patients, 86.2% rated the procedure as very good or good, and pain relief was observed already after 1 month.⁸⁴ However, there have been reported complications associated with this procedure. Durdag and colleagues reoperated a woman with a GelStix® implanted as she was admitted with severe radicular pain.¹⁴⁹ The pain was linked to a fragment of implant that had penetrated through an annual tear and caused compression to the spinal root. The authors speculate that the implant may have been initially wrongly placed in the AF, highlighting the importance of the correct placement of the implant.

HA with a similar composition to the solutions used for VS has been used for treatment of the NP. In a 24-week follow-up period, Mazza and co-workers observed relief from chronic lower back pain due to DDD compared to the baseline.⁸⁵ They had two patients drop out due to adverse events, but this is not believed to be related to the treatment. However, the clinical efficacy is proven only over a short time period. Considering the surgical risk related to bypassing vital organs during injection, this therapy can come short when evaluating

it using a cost-benefit analysis. Hence a longer-lasting therapy should be investigated.

Two other solutions have been tried clinically, but seem to have been discontinued. The NuCore® gel for NP replacement consists of elastin and silk co-polymers that are crosslinked in situ.⁸⁹ A 2-year follow-up pilot clinical study with 14 patients demonstrated a significant reduction in back and leg pain, regained disc height, and no side effects.⁸⁹ There have not been any clinical publications on this product since 2009, and it seems to have been discontinued by the supplier. CryoLife started clinical trials on their product BioDisc but have not published the outcome of the trial. In a conference abstract containing interim results, they reported at the 6-month follow-up a decrease in mean Oswestry Disability index from 49.2 to 11 and a decrease in numerical pain score from 5.86 to 1.62,⁸⁸ which could seem promising. However, they also reported that 2 of the 10 patients enrolled experienced recurrent herniation requiring surgery. After this abstract from 2008, there has been no publication, and the product seems to have been discontinued.

Since neither CryoLife nor Spine Wave has disclosed why their products were discontinued, it is not feasible to conclude why they failed to perform in the clinic.

## 4 | REGULATORY CLASSIFICATION AND CONSEQUENCES

From a regulatory perspective, the first step of translating a medical device is to assign it to the appropriate risk classification group, namely risk class. In Europe with the new MDR, this is reasonably straightforward with injectable hydrogels. Because they are implanted, hence in contact with human tissue over a prolonged period and have a biological effect, it becomes a class III device (highest risk level). This means a premarket clinical investigation is mandatory. This can be mitigated if equivalence to a predicate device can be demonstrated. Nevertheless, appropriate equivalence is practically infeasible unless the manufacturer of the new device either (a) also manufactures the predicate device or (b) has a contractual agreement with the manufacturer of the predicate to access all technical information. In the United States, the risk classification differs from Europe as it depends on product device groups. VS products or dental biologics (e.g., Emdogain®) are class III (highest risk), meanwhile more conventional bone graft materials without human growth factors such as Anika's Tactoset® or the DBM solutions are class II. For class II and some class III product groups the 510(k)-pathway can be used if it demonstrates substantial equivalence with existing approved devices, demonstrating that the device is safe and efficient, which is significantly cheaper than introducing a new device. In the case of class III, the 510(k) allows the manufacturer to partially bypass the premarket approval application, meaning they do not need to run a clinical investigation, but this is not applicable for the VS products or dental biological materials discussed here. When the 510(k) is not applicable for the class III devices, the product needs to be evaluated on a case-by-case basis by the authorities (US-FDA).
<table><caption>TABLE 2 List of hydrogel solutions for musculoskeletal therapies regulated as medicinal products</caption><thead><tr><th>Indication and treatment mode</th><th>Product (producer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>FDA approvals</th></tr></thead><tbody><tr><td>Osteoarthritis</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Cartilage regeneration</td><td>NovoCart<br>Inject<sup>150,151</sup><br>(Tetec AG)</td><td>Maleimide functionalized human albumin and HA crosslinked with bisthio-PEG, and autologous chondrocytes</td><td>Arthroscopic injectable autologous chondrocyte transplant</td><td>Needle injection through two-chamber solution allowing in situ polymerization</td><td>FDA phase III trials</td></tr><tr><td>DDD</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Nucleus pulposus replacement</td><td>NovoCart Disc<sup>152,153</sup><br>(Tetec AG)</td><td>As above</td><td>As above</td><td>As above</td><td>FDA phase II trials</td></tr></tbody></table>

Abbreviations: DDD, degenerative disc disease; FDA, Food and Drug Administration; HA, hyaluronic acid; PEG, polyethylene glycol.

In the review, we have focused on discussing hydrogels as medical devices. However, they can also be classified as medicinal products if their main mechanism of action is through pharmacological, metabolic, or immunological means⁶; this would lead them to the so-called “drug approval process.” A couple of hydrogels that are used for the above-described musculoskeletal treatments are classified by the European Medical Agency and the FDA as medicinal products (biologics/drugs) instead of medical devices as they have the characteristics of combinatory products, Advanced Therapy Medicinal Products (ATMP). A summary of these can be found in Table 2.

Over the last couple of decades, there has been a drastic change in the design rationale of orthopedic biomaterials. From passive structures designed for minimal interaction with the surrounding tissue, for example, titanium-based hip implants, the current generation of biomaterials is designed to actively interact with the surrounding tissue, such as scaffolds for tissue regeneration that stimulates tissue growth. This means that the product's mechanism of action starts approaching that of medicinal products, which will change the applicable regulation framework.¹⁵⁴ Hence engineers need to carefully consider regulatory classification when designing hydrogels. If a hydrogel solution is classified as a medicinal product, it increases the documentation and overall market entry requirement and requires larger and more costly clinical trials. Compared to medical devices, the therapy will take significantly longer time for clinical translation, the R&D investment costs will increase drastically, and the product will eventually be sold at a higher price to the healthcare providers. Moreover, there will be longer product cycles, which means less innovation. In the United States, it takes on average 12 years from pre-clinical trials to market approval for drugs while it only takes 3–7 years for medical devices, and the development costs will increase from the range of tens of millions of dollars for medical devices up to the excess of $1 billion for pure drugs.¹⁵⁵,¹⁵⁶

Products where a medical device (i.e., the gel) carries a therapeutic agent such as growth factor or expanded cells no longer gets its primary mode of action through physical means, and the classification changes to medicinal products. For example, Tetac AG (Germany) has developed two such products for cartilage treatment (NovoCart Inject®) and IVD regeneration (NovoCart Disc®). The NovoCart® gel

functions as an autologous chondrocyte carrier and is used in a 2-step surgical procedure, hence is regulated after the complex ATMP framework. In the first step, the chondrocytes are harvested and expanded in GMP facilities. In the second step, the cells are added to a liquid consisting of human albumin and HA. During the injection procedure, the cell/polymer mixture is mixed with a bisthio-PEG crosslinker which causes the crosslinking of the gel in situ through a Michael-type addition reaction between the thiol groups of the PEG and the maleimide groups of the functionalized human albumin.¹⁵⁷ In a short-term follow-up (12 months) for cartilage regeneration, they could observe a reduction in pain, an increase in activity and quality of life among the patients.¹⁵⁰ In a smaller 24-months study, they demonstrated clinically favorable outcome in terms of reduced pain and a MOCART 2.0 score of 70 ± 13.6, suggesting cartilage regrowth with morphological integrity.¹⁵¹ The MOCART 2.0 scoring system uses MRI to quantify the quality of cartilage repair tissue by giving it a score between 0 (worst) to 100 (best).¹⁵⁸ The Novocart® inject solution has also been tried clinically for NP regeneration. So far, the phase I part of the joint I/II trials have not raised any concerns about the safety of the product.¹⁵²

## 5 | FROM LAB TO CLINIC AND EMERGENCE OF POST-MARKET SURVEILLANCE

Translating hydrogels as medical devices is a time-consuming process, and care should be taken to have a clear plan from design to pre-clinical and clinical investigation. The steps from hydrogel development to clinical approval and post-market surveillance have been illustrated in Figure 5. First, the hydrogel needs to be developed; the details of this process have been described above. A thorough material characterization is mandatory for scientific and clinical perspectives and is also useful when explaining the mechanism of action to the notifying body or the US-FDA. Thereafter, it is mandatory to demonstrate biocompatibility according to the applicable ISO 10993 standards, where the manufacturer must justify which are applicable and which are not. A natural sequence for hydrogels for musculoskeletal
<table><caption>TABLE 2 List of hydrogel solutions for musculoskeletal therapies regulated as medicinal products</caption><thead><tr><th>Indication and treatment mode</th><th>Product (producer)</th><th>Composition</th><th>Delivery method</th><th>Therapeutic claim</th><th>FDA approvals</th></tr></thead><tbody><tr><td>Osteoarthritis</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Cartilage regeneration</td><td>NovoCart<br>Inject<sup>150,151</sup><br>(Tetec AG)</td><td>Maleimide functionalized human albumin and HA crosslinked with bisthio-PEG, and autologous chondrocytes</td><td>Arthroscopic injectable autologous chondrocyte transplant</td><td>Needle injection through two-chamber solution allowing in situ polymerization</td><td>FDA phase III trials</td></tr><tr><td>DDD</td><td></td><td></td><td></td><td></td><td></td></tr><tr><td>Nucleus pulposus replacement</td><td>NovoCart Disc<sup>152,153</sup><br>(Tetec AG)</td><td>As above</td><td>As above</td><td>As above</td><td>FDA phase II trials</td></tr></tbody></table>

Abbreviations: DDD, degenerative disc disease; FDA, Food and Drug Administration; HA, hyaluronic acid; PEG, polyethylene glycol.

In the review, we have focused on discussing hydrogels as medical devices. However, they can also be classified as medicinal products if their main mechanism of action is through pharmacological, metabolic, or immunological means⁶; this would lead them to the so-called “drug approval process.” A couple of hydrogels that are used for the above-described musculoskeletal treatments are classified by the European Medical Agency and the FDA as medicinal products (biologics/drugs) instead of medical devices as they have the characteristics of combinatory products, Advanced Therapy Medicinal Products (ATMP). A summary of these can be found in Table 2.

Over the last couple of decades, there has been a drastic change in the design rationale of orthopedic biomaterials. From passive structures designed for minimal interaction with the surrounding tissue, for example, titanium-based hip implants, the current generation of biomaterials is designed to actively interact with the surrounding tissue, such as scaffolds for tissue regeneration that stimulates tissue growth. This means that the product's mechanism of action starts approaching that of medicinal products, which will change the applicable regulation framework.¹⁵⁴ Hence engineers need to carefully consider regulatory classification when designing hydrogels. If a hydrogel solution is classified as a medicinal product, it increases the documentation and overall market entry requirement and requires larger and more costly clinical trials. Compared to medical devices, the therapy will take significantly longer time for clinical translation, the R&D investment costs will increase drastically, and the product will eventually be sold at a higher price to the healthcare providers. Moreover, there will be longer product cycles, which means less innovation. In the United States, it takes on average 12 years from pre-clinical trials to market approval for drugs while it only takes 3–7 years for medical devices, and the development costs will increase from the range of tens of millions of dollars for medical devices up to the excess of $1 billion for pure drugs.¹⁵⁵,¹⁵⁶

Products where a medical device (i.e., the gel) carries a therapeutic agent such as growth factor or expanded cells no longer gets its primary mode of action through physical means, and the classification changes to medicinal products. For example, Tetac AG (Germany) has developed two such products for cartilage treatment (NovoCart Inject®) and IVD regeneration (NovoCart Disc®). The NovoCart® gel

functions as an autologous chondrocyte carrier and is used in a 2-step surgical procedure, hence is regulated after the complex ATMP framework. In the first step, the chondrocytes are harvested and expanded in GMP facilities. In the second step, the cells are added to a liquid consisting of human albumin and HA. During the injection procedure, the cell/polymer mixture is mixed with a bisthio-PEG crosslinker which causes the crosslinking of the gel in situ through a Michael-type addition reaction between the thiol groups of the PEG and the maleimide groups of the functionalized human albumin.¹⁵⁷ In a short-term follow-up (12 months) for cartilage regeneration, they could observe a reduction in pain, an increase in activity and quality of life among the patients.¹⁵⁰ In a smaller 24-months study, they demonstrated clinically favorable outcome in terms of reduced pain and a MOCART 2.0 score of 70 ± 13.6, suggesting cartilage regrowth with morphological integrity.¹⁵¹ The MOCART 2.0 scoring system uses MRI to quantify the quality of cartilage repair tissue by giving it a score between 0 (worst) to 100 (best).¹⁵⁸ The Novocart® inject solution has also been tried clinically for NP regeneration. So far, the phase I part of the joint I/II trials have not raised any concerns about the safety of the product.¹⁵²

## 5 | FROM LAB TO CLINIC AND EMERGENCE OF POST-MARKET SURVEILLANCE

Translating hydrogels as medical devices is a time-consuming process, and care should be taken to have a clear plan from design to pre-clinical and clinical investigation. The steps from hydrogel development to clinical approval and post-market surveillance have been illustrated in Figure 5. First, the hydrogel needs to be developed; the details of this process have been described above. A thorough material characterization is mandatory for scientific and clinical perspectives and is also useful when explaining the mechanism of action to the notifying body or the US-FDA. Thereafter, it is mandatory to demonstrate biocompatibility according to the applicable ISO 10993 standards, where the manufacturer must justify which are applicable and which are not. A natural sequence for hydrogels for musculoskeletal
<figure><img src="image_7.png" /><figcaption>**FIGURE 5** (a) Schematic illustrating the main stages involved in the clinical translation of injectable hydrogels. (b) Suggested framework for industrial-academic collaboration on post-market clinical follow-up (PMCF) in accordance with the requirement of the EU MDR 2017/745 regulations for Class III (high-risk) medical devices such as hydrogels. The responsibility division is not resolute, with the expectation of the data collection and the annual report, and should be delegated on a case-to-case basis. FDA, US Food and Drug Administration; ISO, International Organization for Standardization; NB, notifying body</figcaption></figure>

application (implantable with long-term tissue contact) is first characterizing the material's chemical properties according to ISO 10993-18:2020, then in vitro cytotoxicity according to ISO 10993-5:2009, and finally pre-clinical trials according to the ISO 10993-6:2016 where both the local and systematic response should be evaluated in a reliable animal model. If these are followed diligently and the animal model is well designed, it should cover most of the documentation requirements of the regulatory body, and most of the other ISO 10993 standards can be considered non-applicable. However, a justification for this must be given in the device's risk management file.

Before the pre-clinical trials, it is important to have a clear idea of the clinical claims that shall be demonstrated in the animal model stage. Indeed, in pre-clinical trials, selecting an animal and implantation site that represents the clinical pathophysiology and loading is essential. This has been discussed in detail for injectable bone

substitutes by Bongio and colleagues.¹⁵⁹ If the pre-clinical trials are successful, it is necessary to go through clinical trials. Since hydrogels tend to be short-term (>60 min contact with tissue) or long-term implants (>30 days contact with tissue) with a biological effect, they will be classified as high-risk (class III) medical devices according to the EU MDR¹⁶⁰ and require thorough documentation on safety and efficacy. The new EU MDR requires a more comprehensive clinical evaluation than the former regulation, focusing on both direct clinical investigation and literature/market analysis. The specific requirements have been discussed from a notifying body's perspective by Holborow.¹⁶¹ It is worth noting that the new regulations require the clinical investigation to have a representative patient group to the EU population, the participation number must be demonstrated statistically to be large enough to be appropriate for demonstrating safety and performance, and the length and follow-up intervals must give a good picture of the lifetime of the
<figure><img src="image_7.png" /><figcaption>**FIGURE 5** (a) Schematic illustrating the main stages involved in the clinical translation of injectable hydrogels. (b) Suggested framework for industrial-academic collaboration on post-market clinical follow-up (PMCF) in accordance with the requirement of the EU MDR 2017/745 regulations for Class III (high-risk) medical devices such as hydrogels. The responsibility division is not resolute, with the expectation of the data collection and the annual report, and should be delegated on a case-to-case basis. FDA, US Food and Drug Administration; ISO, International Organization for Standardization; NB, notifying body</figcaption></figure>

application (implantable with long-term tissue contact) is first characterizing the material's chemical properties according to ISO 10993-18:2020, then in vitro cytotoxicity according to ISO 10993-5:2009, and finally pre-clinical trials according to the ISO 10993-6:2016 where both the local and systematic response should be evaluated in a reliable animal model. If these are followed diligently and the animal model is well designed, it should cover most of the documentation requirements of the regulatory body, and most of the other ISO 10993 standards can be considered non-applicable. However, a justification for this must be given in the device's risk management file.

Before the pre-clinical trials, it is important to have a clear idea of the clinical claims that shall be demonstrated in the animal model stage. Indeed, in pre-clinical trials, selecting an animal and implantation site that represents the clinical pathophysiology and loading is essential. This has been discussed in detail for injectable bone

substitutes by Bongio and colleagues.¹⁵⁹ If the pre-clinical trials are successful, it is necessary to go through clinical trials. Since hydrogels tend to be short-term (>60 min contact with tissue) or long-term implants (>30 days contact with tissue) with a biological effect, they will be classified as high-risk (class III) medical devices according to the EU MDR¹⁶⁰ and require thorough documentation on safety and efficacy. The new EU MDR requires a more comprehensive clinical evaluation than the former regulation, focusing on both direct clinical investigation and literature/market analysis. The specific requirements have been discussed from a notifying body's perspective by Holborow.¹⁶¹ It is worth noting that the new regulations require the clinical investigation to have a representative patient group to the EU population, the participation number must be demonstrated statistically to be large enough to be appropriate for demonstrating safety and performance, and the length and follow-up intervals must give a good picture of the lifetime of the
device.¹⁶² This means that it is technically enough to conduct one clinical trial to get CE approval, but it must be large enough to be exhaustively representative. It is also a requirement that the clinical trials must strictly follow Good Clinical Practice (GCP) guidelines and ISO 14155:2020; hence they must be approved by an ethical committee set up according to national law in the EU member state where the clinical trials are conducted.¹⁶²

The clinical trials for medical devices differ from medicinal prod-
ucts, where there are distinct phases in the clinical trials. The typical
set of clinical trials for drugs consists of phase I where safety is dem-
onstrated on a small number of healthy participants, phase II where
efficacy is demonstrated on a moderate number of participants, and
phase III where efficacy is demonstrated on a larger number of partici-
pants. The phase III trial, which ideally is double-blind and randomized,
can involve up to thousands of participants lasting months or years.¹⁶³
This might not be feasible or ethical for medical devices. For example,
although saline solution as a control for VS is standard procedure, a
sham control for an orthopedic bone graft could do significant damage
to the patient and thereby be unethical. For these risk cases, using the
current treatment alternative as a positive control could be a good
alternative, such as autografts as a control for bone graft substitutes.
This allows the manufacturers to benchmark their technology, and it
is easier to demonstrate its clinical claims and the value provided to
patients and healthcare providers. Since clinical trials directly affect
the patient's health, patient safety and ethical standards should be
central in clinical trials to reassure a high-quality standard. The Decla-
ration of Helsinki is an excellent guideline for meeting the ethical stan-
dards, together with GCP and ISO 14155:2020.

Notably, the new EU MDR requires a post-market surveillance
register for medical devices (EUDAMED). This is inspired by the suc-
cessful implementation of orthopedic device registries and the quality
of the data these have provided.⁴ With this registry, the regulation
requires continuous data gathering and analysis. More specifically the
MDR article 83 states¹⁶²: “The post-market surveillance system shall be
suited to actively and systematically gathering, recording and analysing
relevant data on the quality, performance and safety of a device through-
out its entire lifetime, and to drawing the necessary conclusions and to
determining, implementing and monitoring any preventive and corrective
actions.” The medical device industry is characterized by a lot of small,
niche suppliers. In Europe, out of 33,000 medical technology compa-
nies, 95% are small or medium enterprises (<250 employees), and a
majority are small or micro-sized companies (<50 employees).¹⁶⁴ The
limited manpower makes it challenging for these companies to desig-
nate and dedicate personnel for the post-market clinical follow-up.
This provides a golden opportunity for academic researchers to col-
laborate with these companies to analyze the clinical data, and aca-
demics can use their understanding of fundamental biological and
clinical mechanisms to explain the collected observations, for example,
evaluating porcine versus bovine gelatin in the bone graft
SmartBone.¹⁶⁵ If the data are published, it will indeed help the wider
research community. Meanwhile, the companies will benefit from this
as they can leverage experienced personnel to analyze and explain
complex data.

6 | CONCLUDING REMARK AND FUTURE DIRECTION

There is a tremendous discrepancy between the intensity of academic research on hydrogels and the number of products that have been clinically translated for the treatment of musculoskeletal defects. When developing hydrogels, it is crucial to consider the clinical potential of the material, and here pre-clinical and clinical trials are key in predicting whether a material candidate will make it past the evaluation of the regulatory body and succeed clinically. On top of that, practical factors such as the cost of the product, scalability, and ease of use in the clinic should be considered at an early point, together with quality assurance and regulatory affairs matters. As demonstrated in this review, the clinically available materials tend to have extensive clinical documentation, but the understanding and documentation of the hydrogel composition tend to be limited. Concurrently, materials that are intensely investigated in academia and have been thoroughly characterized physiochemically and in vitro are not the ones that have made it to the clinic. When searching “GelMA” on PubMed, it yields 701 articles from the last 11 years. Of these, none are clinical trials. The fact that GelMA has not made it to the clinic is likely a consequence that regulatory bodies are primarily concerned about the material’s clinical history. Hence, materials that have made it to the clinic before increased documentation requirements are favorable to use in new implants. Meanwhile, new biomaterials are now expensive and scientifically challenging to translate. It can also indicate that academic research environments need to invest more resources to mature the technology through in vitro, pre-clinical and clinical trials. Particularly, a comprehensive characterization of physiochemical properties, in vitro testing, and use of advanced characterization in animal trials will be helpful for industry, both because it helps them understand the potential of the biomaterial and because it can assist in explaining a device’s mechanism of action. A complete understanding of a device’s mechanism of action is essential for approval under the new MDR. To increase the likelihood of industrial adoption academics should also demonstrate that any new therapeutic agents can withstand appropriate manufacturing, for instance how an osteoinductive peptide can withstand manufacturing processes with DCM and other solvents,¹⁶⁶,¹⁶⁷ and sterilization processes (e.g., autoclaving, gamma/beta-irradiation [typically 25 kGy], ethylene oxide) without compromising its efficacy. Additionally, verify that the clinical effect and sterility can be maintained with storage over an extended time period in accordance with ISO 11737-2:2020.

In vitro testing is very important for understanding isolated mech-
anisms. However, in our experience¹⁶⁷⁻¹⁶⁹ there are major differences
in response to biomaterials during in vitro tests, where single cell
types are used, and in vivo, where there is an assortment of cell types
interacting.³³,¹⁷⁰ Although there is progress in technology such as
organ-on-chip¹⁷¹,¹⁷² or co-cultures,¹⁷³ they are yet not capable of
mimicking the complexity of tissue response to biomaterials. Simulta-
neously, animal trials should be kept to a minimum for ethical and eco-
nomic reasons. To obtain adequate documentation and keep animal
trials to a minimum, care should be taken in acquiring high quality
device.¹⁶² This means that it is technically enough to conduct one clinical trial to get CE approval, but it must be large enough to be exhaustively representative. It is also a requirement that the clinical trials must strictly follow Good Clinical Practice (GCP) guidelines and ISO 14155:2020; hence they must be approved by an ethical committee set up according to national law in the EU member state where the clinical trials are conducted.¹⁶²

The clinical trials for medical devices differ from medicinal prod-
ucts, where there are distinct phases in the clinical trials. The typical
set of clinical trials for drugs consists of phase I where safety is dem-
onstrated on a small number of healthy participants, phase II where
efficacy is demonstrated on a moderate number of participants, and
phase III where efficacy is demonstrated on a larger number of partici-
pants. The phase III trial, which ideally is double-blind and randomized,
can involve up to thousands of participants lasting months or years.¹⁶³
This might not be feasible or ethical for medical devices. For example,
although saline solution as a control for VS is standard procedure, a
sham control for an orthopedic bone graft could do significant damage
to the patient and thereby be unethical. For these risk cases, using the
current treatment alternative as a positive control could be a good
alternative, such as autografts as a control for bone graft substitutes.
This allows the manufacturers to benchmark their technology, and it
is easier to demonstrate its clinical claims and the value provided to
patients and healthcare providers. Since clinical trials directly affect
the patient's health, patient safety and ethical standards should be
central in clinical trials to reassure a high-quality standard. The Decla-
ration of Helsinki is an excellent guideline for meeting the ethical stan-
dards, together with GCP and ISO 14155:2020.

Notably, the new EU MDR requires a post-market surveillance
register for medical devices (EUDAMED). This is inspired by the suc-
cessful implementation of orthopedic device registries and the quality
of the data these have provided.⁴ With this registry, the regulation
requires continuous data gathering and analysis. More specifically the
MDR article 83 states¹⁶²: “The post-market surveillance system shall be
suited to actively and systematically gathering, recording and analysing
relevant data on the quality, performance and safety of a device through-
out its entire lifetime, and to drawing the necessary conclusions and to
determining, implementing and monitoring any preventive and corrective
actions.” The medical device industry is characterized by a lot of small,
niche suppliers. In Europe, out of 33,000 medical technology compa-
nies, 95% are small or medium enterprises (<250 employees), and a
majority are small or micro-sized companies (<50 employees).¹⁶⁴ The
limited manpower makes it challenging for these companies to desig-
nate and dedicate personnel for the post-market clinical follow-up.
This provides a golden opportunity for academic researchers to col-
laborate with these companies to analyze the clinical data, and aca-
demics can use their understanding of fundamental biological and
clinical mechanisms to explain the collected observations, for example,
evaluating porcine versus bovine gelatin in the bone graft
SmartBone.¹⁶⁵ If the data are published, it will indeed help the wider
research community. Meanwhile, the companies will benefit from this
as they can leverage experienced personnel to analyze and explain
complex data.

6 | CONCLUDING REMARK AND FUTURE DIRECTION

There is a tremendous discrepancy between the intensity of academic research on hydrogels and the number of products that have been clinically translated for the treatment of musculoskeletal defects. When developing hydrogels, it is crucial to consider the clinical potential of the material, and here pre-clinical and clinical trials are key in predicting whether a material candidate will make it past the evaluation of the regulatory body and succeed clinically. On top of that, practical factors such as the cost of the product, scalability, and ease of use in the clinic should be considered at an early point, together with quality assurance and regulatory affairs matters. As demonstrated in this review, the clinically available materials tend to have extensive clinical documentation, but the understanding and documentation of the hydrogel composition tend to be limited. Concurrently, materials that are intensely investigated in academia and have been thoroughly characterized physiochemically and in vitro are not the ones that have made it to the clinic. When searching “GelMA” on PubMed, it yields 701 articles from the last 11 years. Of these, none are clinical trials. The fact that GelMA has not made it to the clinic is likely a consequence that regulatory bodies are primarily concerned about the material’s clinical history. Hence, materials that have made it to the clinic before increased documentation requirements are favorable to use in new implants. Meanwhile, new biomaterials are now expensive and scientifically challenging to translate. It can also indicate that academic research environments need to invest more resources to mature the technology through in vitro, pre-clinical and clinical trials. Particularly, a comprehensive characterization of physiochemical properties, in vitro testing, and use of advanced characterization in animal trials will be helpful for industry, both because it helps them understand the potential of the biomaterial and because it can assist in explaining a device’s mechanism of action. A complete understanding of a device’s mechanism of action is essential for approval under the new MDR. To increase the likelihood of industrial adoption academics should also demonstrate that any new therapeutic agents can withstand appropriate manufacturing, for instance how an osteoinductive peptide can withstand manufacturing processes with DCM and other solvents,¹⁶⁶,¹⁶⁷ and sterilization processes (e.g., autoclaving, gamma/beta-irradiation [typically 25 kGy], ethylene oxide) without compromising its efficacy. Additionally, verify that the clinical effect and sterility can be maintained with storage over an extended time period in accordance with ISO 11737-2:2020.

In vitro testing is very important for understanding isolated mech-
anisms. However, in our experience¹⁶⁷⁻¹⁶⁹ there are major differences
in response to biomaterials during in vitro tests, where single cell
types are used, and in vivo, where there is an assortment of cell types
interacting.³³,¹⁷⁰ Although there is progress in technology such as
organ-on-chip¹⁷¹,¹⁷² or co-cultures,¹⁷³ they are yet not capable of
mimicking the complexity of tissue response to biomaterials. Simulta-
neously, animal trials should be kept to a minimum for ethical and eco-
nomic reasons. To obtain adequate documentation and keep animal
trials to a minimum, care should be taken in acquiring high quality
in vivo data. The ISO 10993-6 (Test for local effect after implantation) requires only local microscopic assessment using histology. Using only this method gives an incomplete picture as conventional histology does not give spatial information or confirm certain biomarkers.¹⁷⁴ Hence, utilizing additional methods such as cone beam computed tomography,¹⁶⁷ microCT (µCT),¹⁶⁷ immunohistochemistry,¹⁶⁷,¹⁷⁵⁻¹⁷⁷ small-angle X-ray scattering,¹⁶⁷,¹⁷⁸ X-Ray diffraction analysis,¹⁶⁷,¹⁷⁸ and more newly developed techniques such as fluorescent labeling of abundant reactive entities,¹⁶⁷ optical photothermal infrared microscopy,¹⁶⁷ and nanoscale atomic force microscopy-infrared¹⁶⁷ can give a comprehensive understanding of the material's mechanism of action. Furthermore, there has been an increased focus on the use of intravital microscopy such as fluorescence lifetime imaging microscopy and Raman spectroscopy as their subcellular resolution (approx. 500 nm) allows for studying in detail in vivo host response to implants and for monitoring of implant biology over time in small animal models.³⁰ If academics bring their material candidates all the way through animal trials and conduct thorough in vivo characterization, it will assist industrial R&D engineers in making an educated choice of biomaterials in their medical device design. Realizing funding limits related to translational research will require the industry to support the financing of these research activities in active collaborations.

## ACKNOWLEDGMENTS

ØØ and HJH were funded by the Research Council Norway through grant no. 327512. ØØ acknowledges additional funding from the Aker Scholarship. JPW and MMS were funded by the grant from the UK Regenerative Medicine Platform “Acellular/Smart Materials—3D Architecture” (MR/R015651/1). CE received funding from the European Union’s Horizon 2020 research and innovation program under the Marie Skłodowska-Curie grant agreement no. 801133. All figures were created with BioRender.com.

## CONFLICT OF INTEREST

M.M.S. is co-inventor on patent application 1911235.8 submitted by Imperial College London, related to ultrasound-triggered liposome payload release.

## AUTHOR CONTRIBUTIONS

**Øystein Øvrebø:** Conceptualization (equal); funding acquisition (equal); visualization (lead); writing – original draft (lead). **Giuseppe Perale:** Conceptualization (equal); supervision (supporting); writing – original draft (supporting). **Jonathan P. Wojciechowski:** Visualization (supporting); writing – review and editing (equal). **Cécile Echalier:** Visualization (supporting); writing – review and editing (equal). **Jonathan R. T. Jeffers:** Supervision (supporting); writing – review and editing (equal). **Molly M. Stevens:** Funding acquisition (equal); supervision (supporting); writing – review and editing (equal). **Håvard J. Haugen:** Conceptualization (equal); funding acquisition (equal); project administration (equal); supervision (equal); writing – original draft (equal). **Filippo Rossi:** Conceptualization (equal); project administration (equal); supervision (equal); writing – original draft (equal).

## PEER REVIEW

The peer review history for this article is available at https://publons.com/publon/10.1002/btm2.10295.

## DATA AVAILABILITY STATEMENT

The PubMed search results on "GelMA" are available upon request from the corresponding author. Otherwise, no data have been used without a reference.

## ORCID

Øystein Øvrebø https://orcid.org/0000-0001-9170-0202

Giuseppe Perale https://orcid.org/0000-0002-6791-3141

Jonathan P. Wojciechowski https://orcid.org/0000-0002-6272-515X

Cécile Echalier https://orcid.org/0000-0002-9733-5477

Jonathan R. T. Jeffers https://orcid.org/0000-0001-8496-6889

Molly M. Stevens https://orcid.org/0000-0002-7335-266X

Håvard J. Haugen https://orcid.org/0000-0002-6690-7233

Filippo Rossi https://orcid.org/0000-0003-2665-120X

## REFERENCES

1. Zhang YS, Khademhosseini A. Advances in engineering hydrogels. *Science*. 2017;356:6337. doi:10.1126/science.aaf3627

2. Tibbitt MW, Anseth KS. Hydrogels as extracellular matrix mimics for 3D cell culture. *Biotechnol Bioeng.* 2009;103(4):655-663. doi:10.1002/bit.22361

3. Correa S, Grosskopf AK, Lopez Hernandez H, et al. Translational applications of hydrogels. *Chem Rev.* 2021;121(18):11385-11457. doi:10.1021/acs.chemrev.0c01177

4. Melvin T, Torre M. New medical device regulations: the regulator's view. *EFORT Open Rev.* 2019;4(6):351-356. doi:10.1302/2058-5241.4.180061

5. Catoira MC, Gonzalez-Payo J, Fusaro L, Ramella M, Boccafoschi F. Natural hydrogels R&D process: technical and regulatory aspects for industrial implementation. *J Mater Sci Mater Med.* 2020;31(8):64. doi:10.1007/s10856-020-06401-w

6. European Commission. MEDICAL DEVICES: Guidance document – Borderline products, drug-delivery products and medical devices incorporating, as an integral part, an ancillary medicinal substance or an ancillary human blood – MEDDEV 2. 1/3 rev 3 derivative. 2015.

7. Spector M. Biomedical materials to meet the challenges of the aging epidemic. *Biomed Mater.* 2018;13(3):030201. doi:10.1088/1748-605X/aab171

8. Milner PE, Parkes M, Puetzer JL, et al. A low friction, biphasic and boundary lubricating hydrogel for cartilage replacement. *Acta Biomater.* 2018;65:102-111. doi:10.1016/j.actbio.2017.11.002

9. Zhang X, Wang J, Jin H, Wang S, Song W. Bioinspired supramolecular lubricating hydrogel induced by shear force. *J Am Chem Soc.* 2018;140(9):3186-3189. doi:10.1021/jacs.7b12886

10. Douglas TEL, Schietse J, Zima A, et al. Novel self-gelling injectable hydrogel/alpha-tricalcium phosphate composites for bone regeneration: physiochemical and microcomputer tomographical characterization. *J Biomed Mater Res A.* 2018;106(3):822-828. doi:10.1002/jbm.a.36277

11. Qi C, Liu J, Jin Y, et al. Photo-crosslinkable, injectable sericin hydrogel as 3D biomimetic extracellular matrix for minimally invasive repairing cartilage. *Biomaterials.* 2018;163:89-104. doi:10.1016/j.biomaterials.2018.02.016

12. Schmocker A, Khoushabi A, Frauchiger DA, et al. A photo-polymerized composite hydrogel and surgical implanting tool for a
in vivo data. The ISO 10993-6 (Test for local effect after implantation) requires only local microscopic assessment using histology. Using only this method gives an incomplete picture as conventional histology does not give spatial information or confirm certain biomarkers.¹⁷⁴ Hence, utilizing additional methods such as cone beam computed tomography,¹⁶⁷ microCT (µCT),¹⁶⁷ immunohistochemistry,¹⁶⁷,¹⁷⁵⁻¹⁷⁷ small-angle X-ray scattering,¹⁶⁷,¹⁷⁸ X-Ray diffraction analysis,¹⁶⁷,¹⁷⁸ and more newly developed techniques such as fluorescent labeling of abundant reactive entities,¹⁶⁷ optical photothermal infrared microscopy,¹⁶⁷ and nanoscale atomic force microscopy-infrared¹⁶⁷ can give a comprehensive understanding of the material's mechanism of action. Furthermore, there has been an increased focus on the use of intravital microscopy such as fluorescence lifetime imaging microscopy and Raman spectroscopy as their subcellular resolution (approx. 500 nm) allows for studying in detail in vivo host response to implants and for monitoring of implant biology over time in small animal models.³⁰ If academics bring their material candidates all the way through animal trials and conduct thorough in vivo characterization, it will assist industrial R&D engineers in making an educated choice of biomaterials in their medical device design. Realizing funding limits related to translational research will require the industry to support the financing of these research activities in active collaborations.

## ACKNOWLEDGMENTS

ØØ and HJH were funded by the Research Council Norway through grant no. 327512. ØØ acknowledges additional funding from the Aker Scholarship. JPW and MMS were funded by the grant from the UK Regenerative Medicine Platform “Acellular/Smart Materials-3D Architecture” (MR/R015651/1). CE received funding from the European Union's Horizon 2020 research and innovation program under the Marie Skłodowska-Curie grant agreement no. 801133. All figures were created with BioRender.com.

## CONFLICT OF INTEREST

M.M.S. is co-inventor on patent application 1911235.8 submitted by Imperial College London, related to ultrasound-triggered liposome payload release.

## AUTHOR CONTRIBUTIONS

**Øystein Øvrebø:** Conceptualization (equal); funding acquisition (equal); visualization (lead); writing – original draft (lead). **Giuseppe Perale:** Conceptualization (equal); supervision (supporting); writing – original draft (supporting). **Jonathan P. Wojciechowski:** Visualization (supporting); writing – review and editing (equal). **Cécile Echalier:** Visualization (supporting); writing – review and editing (equal). **Jonathan R. T. Jeffers:** Supervision (supporting); writing – review and editing (equal). **Molly M. Stevens:** Funding acquisition (equal); supervision (supporting); writing – review and editing (equal). **Håvard J. Haugen:** Conceptualization (equal); funding acquisition (equal); project administration (equal); supervision (equal); writing – original draft (equal). **Filippo Rossi:** Conceptualization (equal); project administration (equal); supervision (equal); writing – original draft (equal).

## PEER REVIEW

The peer review history for this article is available at https://publons.com/publon/10.1002/btm2.10295.

## DATA AVAILABILITY STATEMENT

The PubMed search results on "GelMA" are available upon request from the corresponding author. Otherwise, no data have been used without a reference.

## ORCID

Øystein Øvrebø https://orcid.org/0000-0001-9170-0202

Giuseppe Perale https://orcid.org/0000-0002-6791-3141

Jonathan P. Wojciechowski https://orcid.org/0000-0002-6272-515X

Cécile Echalier https://orcid.org/0000-0002-9733-5477

Jonathan R. T. Jeffers https://orcid.org/0000-0001-8496-6889

Molly M. Stevens https://orcid.org/0000-0002-7335-266X

Håvard J. Haugen https://orcid.org/0000-0002-6690-7233

Filippo Rossi https://orcid.org/0000-0003-2665-120X

## REFERENCES

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2. Tibbitt MW, Anseth KS. Hydrogels as extracellular matrix mimics for 3D cell culture. *Biotechnol Bioeng*. 2009;103(4):655-663. doi:10.1002/bit.22361

3. Correa S, Grosskopf AK, Lopez Hernandez H, et al. Translational applications of hydrogels. *Chem Rev*. 2021;121(18):11385-11457. doi:10.1021/acs.chemrev.0c01177

4. Melvin T, Torre M. New medical device regulations: the regulator's view. *EFORT Open Rev*. 2019;4(6):351-356. doi:10.1302/2058-5241.4.180061

5. Catoira MC, Gonzalez-Payo J, Fusaro L, Ramella M, Boccafoschi F. Natural hydrogels R&D process: technical and regulatory aspects for industrial implementation. *J Mater Sci Mater Med*. 2020;31(8):64. doi:10.1007/s10856-020-06401-w

6. European Commission. MEDICAL DEVICES: Guidance document – Borderline products, drug-delivery products and medical devices incorporating, as an integral part, an ancillary medicinal substance or an ancillary human blood – MEDDEV 2. 1/3 rev 3 derivative. 2015.

7. Spector M. Biomedical materials to meet the challenges of the aging epidemic. *Biomed Mater*. 2018;13(3):030201. doi:10.1088/1748-605X/aab171

8. Milner PE, Parkes M, Puetzer JL, et al. A low friction, biphasic and boundary lubricating hydrogel for cartilage replacement. *Acta Biomater*. 2018;65:102-111. doi:10.1016/j.actbio.2017.11.002

9. Zhang X, Wang J, Jin H, Wang S, Song W. Bioinspired supramolecular lubricating hydrogel induced by shear force. *J Am Chem Soc*. 2018;140(9):3186-3189. doi:10.1021/jacs.7b12886

10. Douglas TEL, Schietse J, Zima A, et al. Novel self-gelling injectable hydrogel/alpha-tricalcium phosphate composites for bone regeneration: physiochemical and microcomputer tomographical characterization. *J Biomed Mater Res A*. 2018;106(3):822-828. doi:10.1002/jbm.a.36277

11. Qi C, Liu J, Jin Y, et al. Photo-crosslinkable, injectable sericin hydrogel as 3D biomimetic extracellular matrix for minimally invasive repairing cartilage. *Biomaterials*. 2018;163:89-104. doi:10.1016/j.biomaterials.2018.02.016

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47. Chahal AS, Schweikle M, Lian AM, Reseland JE, Haugen HJ, Tiainen H. Osteogenic potential of poly(ethylene glycol)-amorphous calcium phosphate composites on human mesenchymal stem cells. *J Tissue Eng*. 2020;11:2041731420926840. doi:10.1177/2041731420926840

48. Appel EA, Tibbitt MW, Webber MJ, Mattix BA, Veiseh O, Langer R. Self-assembled hydrogels utilizing polymer-nanoparticle interactions. *Nat Commun*. 2015;6(1):6295. doi:10.1038/ncomms7295

49. Wang Q, Mynar JL, Yoshida M, et al. High-water-content mouldable hydrogels by mixing clay and a dendritic molecular binder. *Nature*. 2010;463(7279):339-343. doi:10.1038/nature08693

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52. O'Toole JE, Eichholz KM, Fessler RG. Surgical site infection rates after minimally invasive spinal surgery. *J Neurosurg Spine.* 2009; 11(4):471-476. doi:10.3171/2009.5.SPINE08633

53. Spector M, Lim TC. Injectable biomaterials: a perspective on the next wave of injectable therapeutics. *Biomed Mater.* 2016;11(1):014110. doi:10.1088/1748-6041/11/1/014110

54. Zhao L, Weir MD, Xu HH. An injectable calcium phosphate-alginate hydrogel-umbilical cord mesenchymal stem cell paste for bone tissue engineering. *Biomaterials.* 2010;31(25):6502-6510. doi:10.1016/j.biomaterials.2010.05.017

55. Ren P, Li J, Zhao L, et al. Dipeptide self-assembled hydrogels with shear-thinning and instantaneous self-healing properties determined by peptide sequences. *ACS Appl Mater Interfaces.* 2020;12(19): 21433-21440. doi:10.1021/acsami.0c03038

56. Han S-S, Yoon HY, Yhee JY, et al. In situ cross-linkable hyaluronic acid hydrogels using copper free click chemistry for cartilage tissue engineering. *Polym Chem.* 2018;9(1):20-27.

57. Nele V, Schutt CE, Wojciechowski JP, et al. Ultrasound-triggered enzymatic gelation. *Adv Mater.* 2020;32(7):e1905914. doi: 10.1002/adma.201905914

58. Scinto SL, Bilodeau DA, Hincapie R, et al. Bioorthogonal chemistry. *Nat Rev Methods Primers.* 2021;1(1):1-23. doi:10.1038/s43586-021-00028-z

59. Kettenberger U, Pioletti D; Google Patents, assignee. Composition for bone regeneration. Patent application US20180264175A1. 2018.

60. Kettenberger U, Luginbuehl V, Procter P, Pioletti DP. In vitro and in vivo investigation of bisphosphonate-loaded hydroxyapatite particles for peri-implant bone augmentation. *J Tissue Eng Regen Med.* 2017;11(7):1974-1985. doi:10.1002/term.2094

61. Strand V, Baraf HSB, Lavin PT, Lim S, Hosokawa H. A multicenter, randomized controlled trial comparing a single intra-articular injection of Gel-200, a new cross-linked formulation of hyaluronic acid, to phosphate buffered saline for treatment of osteoarthritis of the knee. *Osteoarthr Cartil.* 2012;20(5):350-356. doi:10.1016/j.joca.2012.01.013

62. Miyamoto K, Sasaki M, Minamisawa Y, Kurahashi Y, Kano H, Ishikawa S. Evaluation of in vivo biocompatibility and biodegradation of photocrosslinked hyaluronate hydrogels (HADgels). *J Biomed Mater Res A.* 2004;70(4):550-559.

63. Brandt KD, Block JA, Michalski JP, Moreland LW, Caldwell JR, Lavin PT. Efficacy and safety of intraarticular sodium hyaluronate in knee osteoarthritis. *ORTHOVISC Study Group. Clin Orthop Relat Res.* 2001;385(385):130-143. doi:10.1097/00003086-200104000-00021

64. Petterson SC, Plancher KD. Single intra-articular injection of lightly cross-linked hyaluronic acid reduces knee pain in symptomatic knee osteoarthritis: a multicenter, double-blind, randomized, placebo-controlled trial. *Knee Surg Sports Traumatol Arthrosc.* 2019;27(6): 1992-2002. doi:10.1007/s00167-018-5114-0

65. Priano F. Early efficacy of intra-articular HYADD(R) 4 (Hymovis(R)) injections for symptomatic knee osteoarthritis. *Joints.* 2017;5(2):79-84. doi:10.1055/s-0037-1603677

66. Bliddal H, Overgaard A, Hartkopp A, Beier J, Conaghan PG, Henriksen M. Polyacrylamide hydrogel injection for knee osteoarthritis: a 6 months prospective study. *J Orthop Res Ther.* 2021;6(2): S278.

67. Bliddal H, Overgaard A, Hartkopp A, Beier J, Conaghan PG, Henriksen M. Polyacrylamide hydrogel injection for knee osteoarthritis: results of a 52 week prospective study. *Osteoarthr Cartil.* 2021;29:S278-S278.

68. Shive MS, Hoemann CD, Restrepo A, et al. BST-CarGel: in situ chondroinduction for cartilage repair. *Oper Tech Orthop.* 2006;16(4): 271-278.

69. Shive MS, Stanish WD, McCormack R, et al. BST-CarGel (R) treatment maintains cartilage repair superiority over micro-fracture at 5 years in a multicenter randomized controlled trial. *Cartilage.* 2015;6(2):62-72. doi:10.1177/1947603514562064

70. Kandel L, Agar G, Elkayam O, et al. A novel approach for knee osteoarthritis using high molecular weight hyaluronic acid conjugated to plasma fibrinogen - interim findings of a double-blind clinical study. *Heliyon.* 2020;6(7):e04475. doi:10.1016/j.heliyon.2020.e04475

71. Schreiner MM, Raudner M, Szomolanyi P, et al. Chondral and osteochondral femoral cartilage lesions treated with GelrinC: significant improvement of radiological outcome over time and zonal variation of the repair tissue based on T2 mapping at 24 months. *Cartilage.* 2020;13:1947603520926702.

72. Trattnig S, Ohel K, Mlynarik V, Juras V, Zbyn S, Korner A. Morphological and compositional monitoring of a new cell-free cartilage repair hydrogel technology - GelrinC by MR using semi-quantitative MOCART scoring and quantitative T2 index and new zonal T2 index calculation. *Osteoarthr Cartil.* 2015;23(12):2224-2232. doi:10.1016/j.joca.2015.07.007

73. Goldshmid R, Cohen S, Shachaf Y, et al. Steric interference of adhesion supports in-vitro chondrogenesis of mesenchymal stem cells on hydrogels for cartilage repair. *Sci Rep.* 2015;5(1):12607. doi: 10.1038/srep12607

74. Heijl L, Heden G, Svardstrom G, Ostgren A. Enamel matrix derivative (EMDOGAIN) in the treatment of intrabony periodontal defects. *J Clin Periodontol.* 1997;24(9 Pt 2):705-714. doi:10.1111/j.1600-051x.1997.tb00253.x

75. McGuire MK, Scheyer ET, Schupbach P. A prospective, case-controlled study evaluating the use of enamel matrix derivative on human buccal recession defects: a human histologic examination. *J Periodontol.* 2016;87(6):645-653. doi:10.1902/jop.2016.150459

76. Wang Z, Lu B, Chen L, Chang J. Evaluation of an osteostimulative putty in the sheep spine. *J Mater Sci Mater Med.* 2011;22(1):185-191. doi:10.1007/s10856-010-4175-5

77. Coathup MJ, Cai Q, Campion C, Buckland T, Blunn GW. The effect of particle size on the osteointegration of injectable silicate-substituted calcium phosphate bone substitute materials. *J Biomed Mater Res B Appl Biomater.* 2013;101(6):902-910. doi:10.1002/jbm.b.32895

78. Coughlan M, Davies M, Mostert AK, et al. A prospective, randomized, multicenter study comparing silicated calcium phosphate versus BMP-2 synthetic bone graft in posterolateral instrumented lumbar fusion for degenerative spinal disorders. *Spine (Phila Pa 1976).* 2018;43(15):E860-E868. doi:10.1097/BRS.00000000000002678

79. Zhang H, Yang L, Yang XG, et al. Demineralized bone matrix carriers and their clinical applications: an overview. *Orthop Surg.* 2019;11(5): 725-737. doi:10.1111/os.12509

80. Cammisa FP Jr, Lowery G, Garfin SR, et al. Two-year fusion rate equivalency between Grafton DBM gel and autograft in posterolateral spine fusion: a prospective controlled trial employing a side-by-side comparison in the same patient. *Spine (Phila Pa 1976).* 2004; 29(6):660-666. doi:10.1097/01.brs.0000116588.17192.b9

81. Martin GJ Jr, Boden SD, Titus L, Scarborough NL. New formulations of demineralized bone matrix as a more effective graft alternative in experimental posterolateral lumbar spine arthrodesis. *Spine (Phila Pa 1976).* 1999;24(7):637-645. doi:10.1097/00007632-199904010-00005

82. Stark M, DeBernardis D, McDowell C, Ford E, McMillan S. Percutaneous skeletal fixation of painful subchondral bone marrow edema utilizing an injectable, synthetic, biocompatible hyaluronic acid-based bone graft substitute. *Arthrosc Tech.* 2020;9(11):e1645-e1650. doi:10.1016/j.eats.2020.07.005

83. Jones JR, Brauer DS, Hupa L, Greenspan DC. Bioglass and bioactive glasses and their impact on healthcare. *Int J Appl Glass Sci.* 2016; 7(4):423-434.
51. Darzi SA, Munz Y. The impact of minimally invasive surgical techniques. *Annu Rev Med.* 2004;55:223-237. doi:10.1146/annurev.med.55.091902.105248

52. O'Toole JE, Eichholz KM, Fessler RG. Surgical site infection rates after minimally invasive spinal surgery. *J Neurosurg Spine.* 2009; 11(4):471-476. doi:10.3171/2009.5.SPINE08633

53. Spector M, Lim TC. Injectable biomaterials: a perspective on the next wave of injectable therapeutics. *Biomed Mater.* 2016;11(1):014110. doi:10.1088/1748-6041/11/1/014110

54. Zhao L, Weir MD, Xu HH. An injectable calcium phosphate-alginate hydrogel-umbilical cord mesenchymal stem cell paste for bone tissue engineering. *Biomaterials.* 2010;31(25):6502-6510. doi:10.1016/j.biomaterials.2010.05.017

55. Ren P, Li J, Zhao L, et al. Dipeptide self-assembled hydrogels with shear-thinning and instantaneous self-healing properties determined by peptide sequences. *ACS Appl Mater Interfaces.* 2020;12(19): 21433-21440. doi:10.1021/acsami.0c03038

56. Han S-S, Yoon HY, Yhee JY, et al. In situ cross-linkable hyaluronic acid hydrogels using copper free click chemistry for cartilage tissue engineering. *Polym Chem.* 2018;9(1):20-27.

57. Nele V, Schutt CE, Wojciechowski JP, et al. Ultrasound-triggered enzymatic gelation. *Adv Mater.* 2020;32(7):e1905914. doi: 10.1002/adma.201905914

58. Scinto SL, Bilodeau DA, Hincapie R, et al. Bioorthogonal chemistry. *Nat Rev Methods Primers.* 2021;1(1):1-23. doi:10.1038/s43586-021-00028-z

59. Kettenberger U, Pioletti D; Google Patents, assignee. Composition for bone regeneration. Patent application US20180264175A1. 2018.

60. Kettenberger U, Luginbuehl V, Procter P, Pioletti DP. In vitro and in vivo investigation of bisphosphonate-loaded hydroxyapatite particles for peri-implant bone augmentation. *J Tissue Eng Regen Med.* 2017;11(7):1974-1985. doi:10.1002/term.2094

61. Strand V, Baraf HSB, Lavin PT, Lim S, Hosokawa H. A multicenter, randomized controlled trial comparing a single intra-articular injection of Gel-200, a new cross-linked formulation of hyaluronic acid, to phosphate buffered saline for treatment of osteoarthritis of the knee. *Osteoarthr Cartil.* 2012;20(5):350-356. doi:10.1016/j.joca.2012.01.013

62. Miyamoto K, Sasaki M, Minamisawa Y, Kurahashi Y, Kano H, Ishikawa S. Evaluation of in vivo biocompatibility and biodegradation of photocrosslinked hyaluronate hydrogels (HADgels). *J Biomed Mater Res A.* 2004;70(4):550-559.

63. Brandt KD, Block JA, Michalski JP, Moreland LW, Caldwell JR, Lavin PT. Efficacy and safety of intraarticular sodium hyaluronate in knee osteoarthritis. *ORTHOVISC Study Group. Clin Orthop Relat Res.* 2001;385(385):130-143. doi:10.1097/00003086-200104000-00021

64. Petterson SC, Plancher KD. Single intra-articular injection of lightly cross-linked hyaluronic acid reduces knee pain in symptomatic knee osteoarthritis: a multicenter, double-blind, randomized, placebo-controlled trial. *Knee Surg Sports Traumatol Arthrosc.* 2019;27(6): 1992-2002. doi:10.1007/s00167-018-5114-0

65. Priano F. Early efficacy of intra-articular HYADD(R) 4 (Hymovis(R)) injections for symptomatic knee osteoarthritis. *Joints.* 2017;5(2):79-84. doi:10.1055/s-0037-1603677

66. Bliddal H, Overgaard A, Hartkopp A, Beier J, Conaghan PG, Henriksen M. Polyacrylamide hydrogel injection for knee osteoarthritis: a 6 months prospective study. *J Orthop Res Ther.* 2021;6(2): S278.

67. Bliddal H, Overgaard A, Hartkopp A, Beier J, Conaghan PG, Henriksen M. Polyacrylamide hydrogel injection for knee osteoarthritis: results of a 52 week prospective study. *Osteoarthr Cartil.* 2021;29:S278-S278.

68. Shive MS, Hoemann CD, Restrepo A, et al. BST-CarGel: in situ chondroinduction for cartilage repair. *Oper Tech Orthop.* 2006;16(4): 271-278.

69. Shive MS, Stanish WD, McCormack R, et al. BST-CarGel (R) treatment maintains cartilage repair superiority over micro-fracture at 5 years in a multicenter randomized controlled trial. *Cartilage.* 2015;6(2):62-72. doi:10.1177/1947603514562064

70. Kandel L, Agar G, Elkayam O, et al. A novel approach for knee osteoarthritis using high molecular weight hyaluronic acid conjugated to plasma fibrinogen - interim findings of a double-blind clinical study. *Heliyon.* 2020;6(7):e04475. doi:10.1016/j.heliyon.2020.e04475

71. Schreiner MM, Raudner M, Szomolanyi P, et al. Chondral and osteochondral femoral cartilage lesions treated with GelrinC: significant improvement of radiological outcome over time and zonal variation of the repair tissue based on T2 mapping at 24 months. *Cartilage.* 2020;13:1947603520926702.

72. Trattnig S, Ohel K, Mlynarik V, Juras V, Zbyn S, Korner A. Morphological and compositional monitoring of a new cell-free cartilage repair hydrogel technology - GelrinC by MR using semi-quantitative MOCART scoring and quantitative T2 index and new zonal T2 index calculation. *Osteoarthr Cartil.* 2015;23(12):2224-2232. doi:10.1016/j.joca.2015.07.007

73. Goldshmid R, Cohen S, Shachaf Y, et al. Steric interference of adhesion supports in-vitro chondrogenesis of mesenchymal stem cells on hydrogels for cartilage repair. *Sci Rep.* 2015;5(1):12607. doi: 10.1038/srep12607

74. Heijl L, Heden G, Svardstrom G, Ostgren A. Enamel matrix derivative (EMDOGAIN) in the treatment of intrabony periodontal defects. *J Clin Periodontol.* 1997;24(9 Pt 2):705-714. doi:10.1111/j.1600-051x.1997.tb00253.x

75. McGuire MK, Scheyer ET, Schupbach P. A prospective, case-controlled study evaluating the use of enamel matrix derivative on human buccal recession defects: a human histologic examination. *J Periodontol.* 2016;87(6):645-653. doi:10.1902/jop.2016.150459

76. Wang Z, Lu B, Chen L, Chang J. Evaluation of an osteostimulative putty in the sheep spine. *J Mater Sci Mater Med.* 2011;22(1):185-191. doi:10.1007/s10856-010-4175-5

77. Coathup MJ, Cai Q, Campion C, Buckland T, Blunn GW. The effect of particle size on the osteointegration of injectable silicate-substituted calcium phosphate bone substitute materials. *J Biomed Mater Res B Appl Biomater.* 2013;101(6):902-910. doi:10.1002/jbm.b.32895

78. Coughlan M, Davies M, Mostert AK, et al. A prospective, randomized, multicenter study comparing silicated calcium phosphate versus BMP-2 synthetic bone graft in posterolateral instrumented lumbar fusion for degenerative spinal disorders. *Spine (Phila Pa 1976).* 2018;43(15):E860-E868. doi:10.1097/BRS.00000000000002678

79. Zhang H, Yang L, Yang XG, et al. Demineralized bone matrix carriers and their clinical applications: an overview. *Orthop Surg.* 2019;11(5): 725-737. doi:10.1111/os.12509

80. Cammisa FP Jr, Lowery G, Garfin SR, et al. Two-year fusion rate equivalency between Grafton DBM gel and autograft in posterolateral spine fusion: a prospective controlled trial employing a side-by-side comparison in the same patient. *Spine (Phila Pa 1976).* 2004; 29(6):660-666. doi:10.1097/01.brs.0000116588.17192.b9

81. Martin GJ Jr, Boden SD, Titus L, Scarborough NL. New formulations of demineralized bone matrix as a more effective graft alternative in experimental posterolateral lumbar spine arthrodesis. *Spine (Phila Pa 1976).* 1999;24(7):637-645. doi:10.1097/00007632-199904010-00005

82. Stark M, DeBernardis D, McDowell C, Ford E, McMillan S. Percutaneous skeletal fixation of painful subchondral bone marrow edema utilizing an injectable, synthetic, biocompatible hyaluronic acid-based bone graft substitute. *Arthrosc Tech.* 2020;9(11):e1645-e1650. doi:10.1016/j.eats.2020.07.005

83. Jones JR, Brauer DS, Hupa L, Greenspan DC. Bioglass and bioactive glasses and their impact on healthcare. *Int J Appl Glass Sci.* 2016; 7(4):423-434.
84. Ceylan A, Asik I, Ozgencil GE, Erken B. Clinical results of intradiscal hydrogel administration (GelStix) in lumbar degenerative disc disease. *Turk J Med Sci.* 2019;49(6):1634-1639. doi:10.3906/sag-1901-1

85. Mazza E, Marcia S, Mondaini F, et al. Efficacy and safety of a novel hydrogel (HYADD4-G) in degenerative disc disease patients: a multicentric open label study. *Eur Rev Med Pharmacol Sci.* 2020;24(5): 2692-2703. doi:10.26355/everrev_202003_20539

86. Mainil-Varlet P, Schiavinato A, Ganster MM. Efficacy evaluation of a new hyaluronan derivative HYADD(R)) 4-G to maintain cartilage integrity in a rabbit model of osteoarthritis. *Cartilage*. 2013;4(1):28-41. doi:10.1177/1947603512455193

87. Lewis G. Nucleus pulposus replacement and regeneration/repair technologies: present status and future prospects. *J Biomed Mater Res B Appl Biomater.* 2012;100(6):1702-1720. doi:10.1002/jbm.b.32712

88. Wardlaw D, Craig NJ, Smith FW, Singal V. Early clinical results of an in situ polymerizing protein hydrogel nuclear repair system. The British Editorial Society of Bone & Joint Surgery; 2008:529.

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