Oxygen-Generating Scaffolds: One Step Closer to the Clinical Translation of Tissue Engineered Products

Abstract

The lack of oxygen supply in engineered constructs has been an ongoing challenge for tissue engineering and regenerative medicine. Upon implantation of an engineered tissue, spontaneous blood vessel formation does not happen rapidly, therefore, there is typically a limited availability of oxygen in engineered biomaterials. Providing oxygen in large tissue-engineered constructs is a major challenge that hinders the development of clinically relevant engineered tissues. Similarly, maintaining adequate oxygen levels in cell-laden tissue engineered products during transportation and storage is another hurdle. There is an unmet demand for functional scaffolds that could actively produce and deliver oxygen, attainable by incorporating oxygen-generating materials. Recent approaches include encapsulation of oxygen-generating agents such as solid peroxides, liquid peroxides, and fluorinated substances in the scaffolds. Recent approaches to mitigate the adverse effects, as well as achieving a sustained and controlled release of oxygen, are discussed. Importance of oxygen-generating materials in various tissue engineering approaches such as ex vivo tissue engineering, in situ tissue engineering, and bioprinting are highlighted in detail. In addition, the existing challenges, possible solutions, and future strategies that aim to design clinically relevant multifunctional oxygen-generating biomaterials are provided in this review paper.

1. Introduction

Organ failure and tissue damage due to various reasons such as accidents, sport injuries, burns, congenital defects, and age-related failures might require a transplant to improve the patient’s condition [1]. In addition, reconstructive surgeries might need multiple types of tissues for rebuilding the defect. In clinical practice, autografts and allografts are commonly used in reconstructive surgeries. Although the current gold standard in the repair of failed tissues is the use of autografts, it is limited by the mismatches in size, formation of a secondary wound at the donor site, and other risks such as such as post-operative infection, and increased patient morbidity [2]. Autografts are limited to tissues such as skin, muscle, cartilage, and bone while the use of allografts can be challenging due to the risks of immune reactions and transmission of diseases. The limited availability of organ donors, challenges with immune rejection, and the potential for disease transmission hampers the success of tissue and organ transplantation [3]. As the gap between the number of prospective organ donors and candidate transplant recipients continues to grow, finding alternative therapies to address this issue is critical. Tissue engineering is an emerging multidisciplinary field that generate engineered 3D tissue constructs and it has been utilized to overcome these challenges [4]. The generation of new tissue constructs can be achieved by developing bioactive cell supportive structures (i.e., scaffolds) with biological signals required for cell growth, proliferation, or differentiation [5]. In addition to biodegradable polymers [6–8], inorganic materials such as hydroxyapatite [9][10], eggshell particles [11] and diatoms [12] have also been used in developing bioactive tissue engineering scaffolds.

Different types of cells are used in tissue engineering, including primary cells isolated from patients, progenitor cells, stem cells, and induced pluripotent stem cells (iPSCs). Stem cells, such as mesenchymal stem cells (MSCs), are commonly used in tissue engineering because of their ability to differentiate into different cells in the presence of the corresponding growth factors and the biomimetic microenvironment [13]. A major advantage of the iPSCs is that they are derived from the donor tissues and can differentiate into a new lineage, making these cells relevant to use with tissue-engineered constructs. In addition to biomaterial scaffolds and cells, the engineered constructs can contain bioactive components to support cell proliferation, differentiation, and vascularization [14]. Some of the cell-supportive factors incorporated into scaffolds are extracellular matrix (ECM)-derived biomolecules, growth factors, mineral components, and nanoparticles [15,16]. In principle, following implantation into the patient, the engineered tissue construct integrates with the existing native tissue and perform the corresponding functions. However, in reality, such conventional tissue engineering approaches suffer from the intrinsic challenges of inhomogeneous proliferation of cells on the scaffolds, failure in vascularization and integration with the host tissue upon implantation, and the risk of rejection of the implanted construct. A relatively new approach, referred as in situ tissue engineering has been developed to overcome some of the challenges associated with ex vivo tissue engineering. In situ tissue engineering involves the direct implantation of cell-free bioactive scaffolds in vivo to facilitate cell migration, differentiation, and repair of the defect. It is important to create a properly functioning microenvironment in the scaffolds to ensure effective cell migration, survival, and regain the functions of the damaged tissues [17,18]. Despite the advantages of ex vivo and in situ tissue engineering approaches in repairing relatively simple tissues, developing engineered constructs that have morphological and structural similarity to the native tissues with complex micron-level architecture is critical for fabricating clinically relevant engineered tissues and organs [19,20]. Recent advancements in additive manufacturing technologies, such as 3D bioprinting, have proved promising for developing complex engineered tissues (e.g., heart, heart valve, kidney, liver) with intricate structures and functions [21,22]. A bioprinted construct containing multiple types of cells arranged in complex structures and hydrogel matrices can help achieve structural and functional similarities with the native tissue.

Regardless of the scaffold fabrication methodologies, oxygen plays a critical role in tissue engineering. Oxygen is essential for cell growth, proliferation, and survival. For example, hyperbaric oxygen therapy (HBOT) enhances cell proliferation and migration, and supports wound healing [23,24]. These effects have inspired researchers to discover new strategies for improving oxygen delivery. It is critical to provide an adequate delivery of oxygen in tissue engineering approaches until the functional vascularization is achieved. Homogeneous distribution of oxygen throughout the 3D of the engineered scaffolds has been a challenge. In relatively thick constructs, diffusion of oxygen is typically limited. Oxygen-generating scaffolds are anticipated to prevent hypoxia-induced cell death and support the clinical translation of cell-laden tissue engineered products. Approaches that can provide controlled and sustained delivery of oxygen within physiological limits in engineered tissue constructs are crucial to ensure their effectiveness. Many studies have investigated the material chemistry and biological aspects of the oxygen-generating agents to defeat the challenges associated with oxygen deficiency in tissue engineering. These studies have used solid peroxides, liquid peroxides, or fluorinated compounds in polymer matrices for the long-term release from the scaffolds. In addition, approaches to reduce the adverse effects of the oxygen generation process, mainly associated with free radical formation, have also been extensively studied. This review discusses the current developments in the use of oxygen-generating biomaterials in different tissue engineering approaches, such as in ex vivo tissue engineering, in situ tissue engineering, and 3D bioprinting, as well as the challenges and opportunities in this research area

2. Current challenges in tissue engineering: Critical role of oxygen

Ideal tissue-engineered scaffolds typically contain the necessary components for the survival and growth of the cells and enable the formation of the tissue of interest [25]. Nutrients and growth factors are essential to engineered tissues; likewise, an optimum oxygen supply is essential. Despite the immense progress in research studies with promising outcomes, clinical translation of tissue engineered products is still challenging. Oxygen is a critical factor required to maintain the cells viable in the engineered tissue constructs. However, distribution of oxygen throughout the scaffolds in 3D is a challenge in tissue engineering. In relatively thin constructs such as 1–2 mm thick scaffolds, the oxygen availability can be improved by allowing the diffusion of media containing dissolved oxygen through the interconnected pores of the scaffold. However, developing highly porous scaffolds with interconnected pores might compromise the mechanical strength, especially for applications that require rigid or stiff scaffolds [26]. This hurdle can significantly hinder the bench-to-bed translation of the tissue engineered products. To address the challenges related to hypoxia in large constructs, scaffolds with interconnected vascular mimetic channels can be generated for media perfusion [27]. Nevertheless, the fabrication of capillaries and functionally robust vascular structures throughout the scaffolds remain to be a major hurdle [28]. Slow vascularization in engineered scaffolds upon implantation is another limitation in the tissue engineering field [29]. Strategies to improve vascularization such as the use of angiogenic molecules, nanomaterials, and gene delivery, might enhance vascularization to some extent. However, utilizing traditional approaches is unlikely to achieve vascularization within a short period of time. Oxygen-generating materials that can provide sustained levels of oxygen over extended periods of time are promising candidates to overcome this hurdle.

To overcome the critical challenges with cell survival, researchers have turned to utilizing cell-free bioactive scaffolds as an in situ tissue engineering approach [30,31]. However, for the success of an in situ tissue engineering approach, the cells in the surrounding tissue would need to migrate to the implantation site and proliferate at an adequate pace to repair and regenerate the defective tissue [32]. Unlike the cell-laden scaffolds, prevascularization or immediate vascularization of the constructs upon implantation is not necessary for this approach. Regardless, simultaneous vascularization with cell migration is needed to prevent the hypoxia-induced cell death. While the approaches that utilize angiogenic growth factors and nanomaterials are promising, angiogenesis is not immediate and does not match the migration rate and proliferation of other cells [33]. Thus, the use of additional oxygen sources in 3D scaffolds might enhance cell survival and support their proliferation until the functional vasculature is developed.

Although 3D bioprinting is a promising approach to generate large tissue constructs, its clinical applications are limited by several factors. Diffusion of media and oxygen throughout the construct as well as the removal of metabolic waste from the bioprinted thick constructs are challenging. In addition to the 3D design strategies to generate constructs with interconnected pores, current bioprinting approaches involve the use of relatively soft hydrogels in bioinks to achieve adequate media diffusion throughout the constructs. However, in mechanically rigid bioprinted constructs, the diffusion of oxygenated media through the entire thickness of the construct can be compromised. Moreover, generating interconnected microscale porosity to facilitate media diffusion is a challenge, mainly due to the limited resolution of traditional printing technologies. Thus, new approaches based on oxygen-generating materials can be developed to supply oxygen rather than relying on the diffusion of the growth media [34]. Moreover, the lack of oxygen upon implantation of the 3D bioprinted scaffolds can result in cell death and ultimate failure of the implanted construct. Therefore, incorporating oxygen-generating materials in bioprinted scaffolds can tackle the challenges related to hypoxia in different stages, such as during the printing process, in vitro maturation, and implantation stages.

Continuous oxygen supply can be affected during the transportation of an engineered construct from the laboratory to the clinical facility. Transporting tissue-engineered products in the frozen metabolically inactive state is not a suitable approach, as potentially harmful substances in the freezing media pose additional challenges [35]. In addition, the viability of the seeded cells in the engineered tissue products can significantly deteriorate when stored under frozen conditions [36]. Materials with high oxygen content might be possibly used to address this limitation when transporting organs or engineered tissues. Hence, incorporating oxygen-generating reagents in 3D scaffolds is a promising approach to increase cell survival for in vitro and in vivo applications as well as clinical evaluation (Fig. 1).

Fig. 1.

The relevance of oxygen-generating materials in tissue engineering. Oxygen-generating materials loaded in tissue engineering scaffolds can provide a controlled and sustained release of oxygen in in vitro and in vivo experimental platforms as well as in clinical interventions. Likewise, providing sustained oxygen generation and minimizing hypoxia in in vitro cultures can support cell proliferation, improve the metabolic activity of cells, and facilitate rapid tissue regeneration. In addition to the benefits obtained during in vitro and in vivo testing, sustained oxygen generation can increase the shelf life of the tissue-engineered products, relieve hypoxic cell death during lengthy implantation procedures, and help integrate the engineered product into the host tissue.

2. Oxygen-Generating Materials for Tissue Engineering

Due to their clinical implications, researchers have studied novel oxygen-generating materials that can provide sustained and controlled release of oxygen to enhance cell survival and function (Camci-Unal et al., 2013b; Oh et al., 2009; Sun et al., 2022). Literature studies have shown that chemical compounds such as magnesium peroxide [40], calcium peroxide (CaO₂) [41–44], sodium percarbonate [45], hydrogen peroxide (aH₂O₂) [46], and fluorinated compounds including perfluoro methyl-cyclohexyl piperidine [47] can generate or provide oxygen in engineered constructs to support tissue repair or regeneration.

An efficient strategy to provide oxygen in engineered tissues is to incorporate oxygen-generating substances such as solid peroxides (CaO₂, magnesium peroxide; MgO₂, and sodium percarbonate; (SPC) 5H₂O₂) that can generate oxygen upon contact with water [39]. While solid peroxides are the most common source of oxygen, liquid H₂O₂ can also be utilized to generate oxygen [48](Fig. 2A). However, CaO₂ and MgO₂ are more sustained sources for in situ oxygen formation than liquid H₂O₂. Therefore, solid peroxides are more commonly employed for oxygen generation for more prolonged periods due to their higher oxygen generation potential [49]. The solid inorganic peroxides have been investigated for oxygen generation in liquid environments [50]. The mechanism at which solid peroxides can generate oxygen and water is when exposed to water, H₂O₂ forms, followed by the decomposition of H₂O₂ into water and oxygen [51] (Fig. 2B). Additionally, perfluorocarbons (PFC) can be used as oxygen sources in tissue engineering applications and can carry large quantities of oxygen (Fig. 2C). Since the density of the PFC droplets is higher than water, they are not miscible in an aqueous solution, leading to oxygen entrapment within PFC bubbles and can transfer oxygen by diffusion in liquid emulsions [52]. Different oxygen-generating materials show varying oxygen release mechanisms and profiles depending upon the chemical properties of the material. A comparative table showing the oxygen release mechanisms and amount of oxygen released from each class has been provided in an earlier article [47].

Fig. 2.

Schematic illustration of the mechanism of oxygen (O₂) generation from various oxygen-generating materials. A) A liquid peroxide, hydrogen peroxide (H₂O₂), loaded polylactide-co-glycolide (PLGA) microparticles in alginate matrix containing catalase can release O₂ by the conversion of H₂O₂ into O₂. B) A widely studied solid peroxide, calcium peroxide (CaO₂) nanoparticles, loaded catalase grafted poly-L-lactic acid (PLLA) solid and hollow microparticles provided O₂ release and supported human mesenchymal stem cell (MSC) proliferation. C) Perfluorocarbon (PFC) emulsions can absorb large quantities of O₂ and release slowly when loaded inside surfactant (sodium dodecyl sulfate, SDS or Tween 20)-based or graphene oxide-based emulsions (PFC@GO). PFC@GO showed slower release compared to surfactant-based emulsions. Fig. A is redrawn from Ref. [54]. Fig. B is adapted with modifications from Ref. [55] with the permission of Elsevier. Fig. C is adapted with modifications from Ref. [56] with the permission of Royal Society of Chemistry (RSC).

Various factors, such as the cell density or cell types in 3D scaffolds, scaffold size, implantation site, and the regeneration time required for the tissue of interest, are carefully evaluated while developing oxygen-generating engineered tissue constructs. For instance, the amount of oxygen required for tissue repair or regeneration depends on the cell type, cell density, and the architecture of engineered tissue. Consequently, a relatively higher local oxygen supply is needed in densely loaded constructs with cells. However, regeneration of less densely populated and more minor tissue defects do not require as much oxygen [53]. Hence, it is crucial to optimize the loading amount and release kinetics of the oxygen-generating materials for specific applications based on the needs of the engineered tissues.

3. Mitigating the Effects of Free Radicals During Oxygen Release

Despite the immense potential of oxygen-generating materials in tissue engineering applications, they can also pose some challenges. One of such challenges is the generation of potentially harmful free radicals. As indicated in equations 1–4 [39], the process takes place in two steps. First, the solid inorganic peroxide undergoes hydrolysis and forms H₂O₂. Subsequently, H₂O₂ transforms into water and oxygen [40,50]. The reaction rate can vary depending on the purity and solubility of the peroxides, presence of antioxidants as well as environmental parameters such as temperature and pH.

$${\text{CaO}}_2(\text{s})+2{\text{H}}_2\text{O}\to \text{Ca}{(\text{OH})}_2(\text{s})+{\text{H}}_2{\text{O}}_2$$ (1)

$${\text{MgO}}_2(\text{s})+2{\text{H}}_2\text{O}\to \text{Mg}{(\text{OH})}_2(\text{s})+{\text{H}}_2{\text{O}}_2$$ (2)

$${[\text{SPC}]}_2.3{\text{H}}_2{\text{O}}_2\to 2{\text{Na}}^{+2}+2{{\text{CO}}_3}^{-}+3{\text{H}}_2{\text{O}}_2$$ (3)

$${\text{H}}_2{\text{O}}_2\to {\text{O}}_2+2{\text{H}}_2\text{O}$$ (4)

Among the different types of solid peroxides, calcium has the highest solubility; therefore, it is more commonly used in developing oxygen-generating scaffolds compared to the sodium or magnesium peroxides (Fig. 3A) [47]. However, the decomposition of CaO₂ happens rapidly upon exposure to water, resulting in a burst release of H₂O₂. Thus, in addition to the approaches to control the decomposition of CaO₂ or its release from the scaffolds, it is crucial to increase the rate of conversion of generated H₂O₂ into O₂. Because having a higher concentration of H₂O₂ than the required levels in tissues can produce adverse biological responses including oxidative stress leading to cell apoptosis [57–59]. Moreover, H₂O₂ generated during the decomposition of solid peroxides can interact with Fe²⁺ ions from the surrounding tissue through the Fenton reaction, producing toxic free hydroxyl radicals (•OH) [60]. Therefore, it is imperative to balance the amount of H₂O₂. Correspondingly, in biological systems, enzyme or non-enzyme-based antioxidants can act as a defense system against free radicals to maintain redox homeostasis [61]. In biological systems, H₂O₂ is formed in a reaction of superoxide anion radical (O₂^•−) with the aid of superoxide dismutase (SOD) (equations 5). To prevent accumulation of H₂O₂ in the cells, catalase converts H₂O₂ into H₂O₂ and O₂ (equation 6) [62].

Fig. 3.

Effect of catalase in mitigating the adverse effects of oxygen-generating materials. A) Schematic illustration of oxygen generation through encapsulated catalase in hydrogels under hypoxia. B) Fluorescent microscopy images showing live/dead cells with different concentrations of catalase after 24 h of cell culture (green: live, red: dead HDF cells). C) Effect of catalase containing PLA microspheres carrying CaO₂ on human adipose derived stem cells (hADSCs). C-i) Viability of hADSCs when cultured with PLA microspheres carrying CaO₂. Catalase containing microspheres and controls are indicated as -G and catalase, respectively. C-ii) SEM images of the microspheres grafted with catalase indicating the adhesion of hASCs. D) Myoblast viability and proliferation upon culturing with GelMA hydrogels containing CaO₂ loaded microparticles. D(i) Fluorescent microscopy images showing live/dead cells encapsulated in GelMA hydrogel, supplemented with three different methods of catalase including control no catalase and pristine GelMA hydrogel. D(ii) Fluorescent microscopy images showing live/dead cells in hydrogels containing three different ratios of CaO₂:catalase (1:3, 1:5 and 1:10 mole:unit catalase)(green: live, red: dead C2C12 cells). Fig. B is reproduced from Ref. [64] with the permission of Elsevier, Fig. C is reproduced from Ref. [55] with the permission of Elsevier. Fig. D is reproduced from Ref. [68] with the permission of Wiley.

$$2{{\text{O}}_2}^{•-}+2{\text{H}}^{+}\underset{SOD}{\to }{\text{H}}_2{\text{O}}_2+{\text{O}}_2$$ (5)

$$2{\text{H}}_2{\text{O}}_2\underset{\mathit{\text{catalase}}}{\to }2{\text{H}}_2\text{O}+{\text{O}}_2$$ (6)

Although the enzyme catalase is innately found in tissues for converting H₂O₂ to O₂, the amount of the enzyme is not sufficient for mitigating the adverse effects of solid peroxides, especially in in vitro cell culture systems. Many tissue engineering studies reported the impact of conversion of H₂O₂ to O₂ in the presence of catalase, whether it is supplemented in the cell culture media or within the scaffold [63,64]. For instance, GelMA-CaO₂ hydrogels were shown to generate oxygen for 5 days under hypoxia in the presence of 100 U/mL catalase in cell culture media [63]. In a recent report, GelMA hydrogels containing CaO₂-PCL microparticles with varying concentrations of CaO₂ and 1 mg/mL catalase in the cell culture media was investigated for bone tissue regeneration [42]. In another work, comparisons were made by using different concentrations of CaO₂ and catalase and their impact on human dermal fibroblast (HDF) cell viability [64]. When the CaO₂ concentration in thiolated gelatin-based hydrogels have increased from 0.25% (w/v) to 0.75% (w/v), there was a significant reduction in cell viability after 24 h culture. Meanwhile, when two different concentrations of catalase (25 and 50 U/mL) were used in the culture media while keeping the CaO₂ concentration constant, the cell viability has increased drastically (Fig. 3B). There are only a few studies that reported the direct grafting of catalase on oxygen-generating materials. In such an attempt, Mohseni-Vadeghani et al., examined the effect of catalase grafting on the surface of oxygen releasing microspheres, especially in improving the cell viability [55]. Briefly, PLA microspheres loaded with CaO₂ were generated by water/oil/water emulsion method. Then, catalase was grafted on the microspheres using ethyl (dimethylaminopropyl) carbodiimide and N-Hydroxysuccinimide method. When human adipose-derived stem cells (hASCs) were cultured with the catalase-grafted microspheres, cells showed significantly higher viability (Fig. 3C–i). Moreover, cells were able to adhere on the catalase-grafted PLA/CaO₂ microspheres (Fig. 3C–ii). In a recent study, Willemen et al. has compared the ability of catalase to decrease H₂O₂ level in GelMA hydrogels loaded with PCL-CaO₂ microparticles [55]. When C2C12 myoblasts were encapsulated in GelMA hydrogels containing 1:5 mole CaO₂: unit catalase, catalase containing hydrogels demonstrated higher viability and elongated cell morphology compared to catalase free groups (Fig. 3D–i). Among hydrogels with 1:3, 1:5, and 1:10 CaO₂− catalase ratio (mole CaO₂: unit catalase), 1:5 ratio shown the highest myoblast activity that was chosen for further studies involving 3D bioprinting (Fig. 3D–ii). Apart from catalase, other antioxidants, e.g., glutathione peroxidase (GPX), curcumin, Vitamin-C, and Vitamin-E, can ensure free radical inactivation [65] [55]. For instance, Vitamin-C was incorporated in a composite hydrogel made of polyacrylamide-sodium alginate and CaO₂ nanoparticles [66]. The study evaluated the differentiation of stem cells under hypoxia. They achieved cell proliferation and differentiation of rat bone mesenchymal stem cells by using Vitamin-C under hypoxic conditions. Another study used antioxidant polyurethane scaffolds made via cryogelation to prevent inactivation of catalase under temperature shifts [41]. They used three concentrations of CaO₂ at 1, 2, 3% (w/v). These scaffolds showed initial burst release of oxygen up to 24.26 ± 0.30 % in the first day of incubation, later demonstrating sustained release. The study reported that 1% CaO₂ (w/v) was sufficient to provide adequate oxygen release to maintain the metabolic activity of the H9c2 cardiomyoblast cells for 10 days under hypoxia. In addition to solid peroxides, the oxides such as manganese (IV) dioxide (MnO₂) and zinc oxide (ZnO) have also been used as H₂O₂ scavengers [67]. For instance, MnO₂ nanoparticles were shown to effectively perform both superoxide (SO) dismutase mimetic and H₂O₂ scavenging activities [67]. In addition, the MnO₂ nanoparticles promoted the proliferation of the MIN6 β-cells and protected them against the damage from H₂O₂.

4. Sustained and Controlled Release of Oxygen from Engineered Scaffolds

Despite the potential toxicity associated with free radicals and H₂O₂ released during oxygen generation process, oxygen itself can be toxic to the cells if released abruptly and excessively [69]. Moreover, the rate at which oxygen is released can influence tissue repair and regeneration. When oxygen is generated too rapidly, supersaturation prevents the oxygen from being utilized by the tissue, leading to hyperoxia-related cytotoxic manifestations (Fig. 4A). On the other hand, if oxygen is released too slowly, it does not deliver sufficient oxygen to the cells to sustain healthy cellular function in the engineered tissue. Thus, it is crucial to control the oxygen release kinetics. While there are different ways to control the release of oxygen from oxygen-generating reagents, polymeric carriers and nanostructured materials are deemed to be the most effective method to release the loaded oxygen in a controlled manner [42,70]. Initial strategies to achieve this goal included using hydrophobic polymer-based microparticles as carriers for oxygen-generating materials [42,71]. Table 1 provides a summary of various studies that used hydrophobic polymer encapsulants for sustained oxygen release from tissue engineered constructs. Utilizing hydrophobic materials to encapsulate oxygen-generating materials decreases the rate of oxygen release due to the slow diffusion of water through the hydrophobic material [53]. On the contrary, hydrophilic carriers encapsulating oxygen-generating materials show rapid water adsorption, driving solid peroxide particles to decompose and produce oxygen more quickly [63]. Several studies have reported the use of relatively stable polymers in physiological conditions, such as PCL [72,73], PLA [74], PLGA [37,75] and other biodegradable polymers for the sustained delivery of oxygen from biomaterial scaffolds. In a recent study, PCL microparticles encapsulated with CaO₂ were developed and presented for potential applications in cardiac [73] and bone [42] tissue engineering applications. When these microparticles were loaded in GelMA-based scaffolds, cell viability and proliferation increased. Interestingly, microparticles loaded with CaO₂ exhibited less toxicity to the cells because of the controlled and sustained release of oxygen for a prolonged period of at least two weeks under hypoxia. A perfluorooctane emulsion oxygen carrier-loaded hollow microparticles (PFO-HPs) released oxygen and prevented cell necrosis for approximately two weeks in a hypoxic environment until new blood vessels formed in the 3D construct [76]. Further, microparticle fabrication techniques like electrospraying can be used for the production of multi-walled sustained oxygen delivery systems [72] (Fig. 4B, C). Cells grown with CaO₂-loaded double-walled microparticles revealed higher viability than those grown with single-walled microparticles (Fig. 4D). Presumably, this was because the double-walled particles were gradually releasing oxygen compared to single-walled particles. In a separate study, researchers developed slow oxygen releasing CaO₂/gelatin microsphere-loaded and 3D printed scaffolds to repair the femoral head damaged by osteonecrosis (ONFH) [43]. The scaffolds containing the bone marrow mesenchymal stem cells (BMSCs) were transplanted into the core depression area in the ONFH rabbit model. The study indicated that CaO₂/gelatin microspheres could continuously release oxygen for 19 days. Further assessments revealed that microspheres in the scaffolds could reduce the local cell apoptosis and enhance the survival of the grafted cells in the host. Such controlled oxygen delivery approaches are highly promising strategies for producing large tissue engineered constructs and possibly clinically relevant organ sized tissue constructs. Moreover, oxygen-generating materials can be designed to tightly modulate the oxygen release kinetics. Studies indicated that due to the lower solubility, MgO₂ can provide the slowest oxygen generation compared to the other solid peroxides. For example, the equilibrium coefficient for MgO₂ and CaO₂ with water are 1.8×10⁻¹¹ and 9.8×10⁻⁹, respectively, suggesting that MgO₂ has lower solubility than CaO₂, therefore, providing a slower reaction rate than CaO₂ [77].

Fig. 4.

Sustained release of oxygen from microparticles. A) Schematic showing the importance of sustained release of oxygen to minimize local hyperoxia and prolong the release. B) Scheme showing the fabrication of bilayered oxygen-generating microparticles by electrospraying technique. C) Representative scanning electron microscopy image of the microparticles. D) Live/dead staining images of MIN6 β cells after culturing with oxygen-generating microparticles for five days (a). Fig. B-D are reproduced with modifications from Ref. [72] with the permission of Wiley.

Table 1.

Tissue engineered constructs using hydrophobic polymers for encapsulation of oxygen-generating reagents to provide sustained oxygen release.

Oxygen-generating material	Hydrophobic polymer used for encapsula tion	Additional materials used	Oxygen release rate from the scaffolds	Application	Outcome	References
CaO2	PCL	Ascorbicacid	3 days	Bone tissue engineering, coating material	CaO2 incorporated electrospun nanofibers demonstrated antibacterial activity against E. coli and S.epidermidis, by reducing approximately 95% and 90% of the colonies after 24 h, respectively.	[129]
H2O2	PLGA, PVP	Catalase, N-isopropyl acrylami de, NIPAAm	About 14 days	Cardiac tissue regeneration	Cardiac differentiation of cardiosphere-derived cells was restored under hypoxic conditions.	[49]
H2O2	PLGA	Catalase, alginate	3 days	Skin wound healing, angiogenesis	Release of oxygen was highest by 24 h, and decreased afterwards until 72 h. No byproducts were observed during the degradation of alginate-based sponge and the scaffold supported the migration and proliferation of HUVEC cells.	[48]
CaO2	PU	-	10 days	Cardiac tissue engineering	Although the oxygen generation was higher in 3% CaO2 containing cryogels then 1% and 2% for 10 days, the H9C2 viability in 2 and %3 CaO2 containing cryogels were lower than 1% starting from day 2.	[41]
CaO2	PDMS	-	4 weeks	Pancreatic tissue engineering	During the first 2 weeks of culture under hypoxia, the disks generated oxygen to a degree that was close to optimal cell culture conditions. Furthermore, The metabolic activity of MIN6 cells under hypoxia doubled in 3 days in presence with CaO2, when compared to control, disks without CaO2.	[53]
CaO2	PLLA	Catalase	15 days	Bone tissue regeneration	Without the presence of catalase, the cell viability was found below 75%. Moreover, hASCs could attach in the microcarrier surfaces, maintained their morphology, and spread well	[55]
CaO2	PCL	Catalase	1 week	General application	Electrosprayed particles with an outer flow rate of 1–4 mL/h showed higher metabolic activity under hypoxia, then cells cultured in PCL microparticles under normoxia.	[72]
CaO2	PCL	Catalase	35 days	General application	As the CaO2 concentration increases (0-40-60 mg/mL), the O2 release was measured from 5 to 29% over the 5 weeks of incubation period.	[73]
H2O2	PLLA, PVP	Catalase	10 days	General application	Particles made from lower polymer concentration showed higher initial burst release and shorter release time (10 days) than polymers with higher concentration (14 days). Moreover, as the particle size is getting bigger from 20 μm to 60 μm, reduced the initial burst release rate.	[74]
CaO2	PLGA	Catalase	14 days	Bone tissue engineering	The particles released between 35–67.5 mmHg oxygen during incubation which allowed the proliferation of mesenchymal stem cells 76% more at the end of two weeks.	[75]
CaO2	PCL, Pluronic F127	-	14 days	Regeneration of blood vessels, vascularization	First 5 days of incubation, hollow particles supported MC3T3-E1 proliferation by increasing the cell number approximately 200% however, after day 5, cell number decreased drastically due to insufficient oxygen generation.	[76]
CaO2	PDMS	-	20 days	Angiogenesis, vascularization	The scaffold OxySite, produced oxygen for 20 days that was sufficient for competent vascularization. 1200 μm thick scaffold 17% of oxygen gradient between the surface and surroundings.	[81]
CaO2	PLA	Catalase	14 days	Bone tissue engineering	Electrospun fibers with two different PLA concentrations (6.5 and 13%) showed that as concentration is increasing the oxygen release was reducing. Moreover, 13% of PLA concentration demonstrated higher calcium deposition by mesenchymal stem cells.	[85]
CaO2	PGS, PCL	Catalase	7 days	Wound healing	Fibers with highest CaO2 concentrations, 10%, demonstrated highest zone of inhibition S.aureus. In addition, the same group showed the uppermost initial oxygen release.	[86]
MnO2	PLGA, collagen	-	1 h	General application	Under two concentrations of H2O2, that mimic oxidative stress, the decomposition of MnO2 was faster for 1 mM than 500 μg/ml H2O2, however cell viability decreased drastically after 1 day of culture with 1 mM while, 500 μg/ml remained showed higher metabolic activity.	[130]
PFC	Methylacr ylamide chitosan	-	8 h	Skin tissue engineering, wound ĥealing	Fibroblasts cultured with with higher fluorine concentrations showed higher metabolic activity. In addition, after the modification of the PFC groups through adding fluorinated aromatic group, oxygen release profile was extended from 8 hours to approximately a week.	[131]
CaO2	PLGA, PVA	Catalase	21 days	General applicaion	First 8 days of release was found higher under hypoxia then control hydrogel in normoxia. However, after day 8, the release from nanoparticles have decreased drastically.	[132]

PCL: polycaprolactone, PLGA: poly(lactide-co-glycolide), PVP: poly(2-vinlypyrridione), PU: polyurethane, PLLA: poly (L-lactic acid), PGS: poly(glycerol sebacate), PDMS: polydimethylsiloxane, PLA: polyvinyl alcohol, PFC: perfluorocarbon, CaO₂: calcium peroxide, MnO₂: manganese dioxide, H₂O₂: hydrogen peroxide.

5. Oxygen-Generating Biomaterials for Ex vivo Tissue Engineering

In ex vivo tissue engineering, scaffolds are combined with cells and biomolecules outside the body to obtain cell-laden tissue constructs for implantation. This approach relies on generating biologically relevant ECM-mimetic constructs in vitro to recapitulate the native tissue architecture, physiology, and biological functions [78]. Furthermore, this strategy involves generating biodegradable porous 3D support materials (scaffolds). Such scaffolds are seeded with patient-derived cells. After sufficient expansion of the cells on the scaffold, the laden tissue constructs can be implanted in the defect site in the patient. However, ex vivo tissue engineering has considerable limitations, such as the complications in constructing highly porous scaffolds, difficulties in cell isolation, time consuming in vitro cell cultures, and the lack of host tissue integration of the engineered tissue. While recent techniques can assemble scaffolds to contain thin fibers with thinner partitions, the mechanical properties of the scaffold can be compromised causing the collapse of the construct. In addition, when the cells are placed in the deep regions of microporous scaffolds, transport of nutrients and oxygen-containing media can be limited to the cells. Thus, limited oxygen diffusion generates hypoxia and eventually induces hypoxia-associated cell death. Providing oxygen using oxygen-generating materials in large 3D scaffolds can address this limitation.

Earlier studies focused on perfusion of oxygen-generating materials along with culture media in vitro. For instance, when oxygen generating PFC emulsions were perfused through capillary channels, engineered cardiac tissue constructs showed higher deoxyribonucleic acid (DNA) content and cardiac specific markers (troponin I, connexin-43). This system also had significantly improved contractile properties as compared to the control constructs [79]. Initial attempts that incorporated oxygen-generating materials in culture media were insufficient to support cell survival upon implantation of the construct in vivo. During implantation of the cell-laden scaffolds in the body, the cells experience hypoxia due to the lack of blood supply. The formation of properly developed vasculature takes several weeks after implantation. The lack of adequately formed vasculature can lead to cell death due to lack of oxygen. To overcome this challenge, attempts were made to include oxygen-generating materials in 3D scaffolds [37]. In a previous work, CaO₂-loaded GelMA hydrogels supplied adequate oxygen, reduced metabolic stress of encapsulated cardiac progenitor cells, and increased the cell viability compared to plain GelMA hydrogels [63]. Although they were able to reduce cell death by limiting hypoxia-induced necrosis, sustained release of oxygen over an extended period was not obtained. Later studies that utilized GelMA hydrogels containing CaO₂-loaded PCL microparticles were able to demonstrate oxygen release for approximately 5 weeks and supported myoblast proliferation [73]. A recent study indicated that pancreatic islets seeded into CaO₂-loaded scaffolds had significantly higher viability and function compared to islets seeded in pristine cryogel-scaffolds [80]. When these islet-seeded scaffolds are transplanted into the epididymal fat pad (EFP) of diabetic mice, the animals re-established glycemic control. The cryogels with 0.25% w/w CaO₂ provided a suitable 3D microenvironment for islet survival, proliferation, and function at the extra-hepatic transplantation site. In another study, CaO₂-loaded polydimethylsiloxane (PDMS) scaffolds, OxySite, provided sufficient local oxygenation for up to 20 days (Fig. 5A, 5B) [81]. Upon implantation of the cell-seeded scaffolds in diabetic Lewis rats, OxySite significantly improved the graft outcomes. Table 2 provides a summary of various approaches used to deliver oxygen from scaffolds in ex vivo tissue engineering.

Fig. 5.

Oxygen-generating materials in ex vivo tissue engineering. A) Schematic illustrating the process of syngeneic rat islet transplantation at the omentum site using an oxygen-generating scaffold. Oxygenation from the OxySite scaffold was hypothesized to reduce the hypoxic period between implantation and the development of functional vasculature, improving the transplantation efficacy. B) Immunohistochemistry staining for insulin, glucagon, and smooth muscle actin within explanted scaffolds. B(i-vi) Images collected from the interface between the scaffold and omentum tissue are designated as the “Outer Region”. B(vii-xii) Images collected from the center of the scaffold are designated as the “Center” (Orange dash line = interface; S = scaffolds; scale bar = 100 μm). B(xiii-xv) Ratio of insulin intensity to DAPI intensity normalized to the corresponding control group. C) Micro-CT scans and the three-dimensional reconstruction of the bone defects 10 weeks after transplantation of oxygen-generating CaO₂/PBMSC composite scaffold. D) Histology of defect site after CaO₂ scaffold/PBMSC transplantation. CaO₂ scaffold/PBMSCs group showed higher bone bridge formation compared to Blank scaffold/PBMSC group. Fig. A and B are reproduced from Ref. [81] with the permission of Elsevier. Fig. C, and D are reproduced with modifications from Ref. [82] with the permission of Elsevier.

Table 2.

Oxygen-generating biomaterials for ex vivo tissue engineering.

Oxygen-generating material	Fabrication method	Scaffold material	Cell type	Other component s	Application	Outcome	Ref.
CaO2	3D printing	PCL/nHA Sodium	BMSCs	alginate, gelatin	Osteonecrosis, angiogenesis	CaO2/gelatin microspheres can release oxygen up to 19 days meanwhile in vivo model on rabbit showed high CD 31 expression 4 weeks after the surgery	[43]
CaO2	Cryogelation	collagen	Islet cells	-	Vascularization	Pancreatic islets seeded into CaO2-loaded scaffolds demonstrated glycemic control after transplanted into epididymal fat pad (EFP) of diabetic mice	[80]
CaO2	Agitation, leaching out	PDMS	Islet cells	-	Angiogenesis, vascularization	The scaffold provided sufficient local oxygenation for up to 20 days. Moreover, the additional oxygen supply provided by the scaffold rescued ~25% more islets	[81]
CaO2	Crosslinkin g with GA/freeze dry	SAP/PLGA	PBMSCs	gelatin	Bone tissue engineering	Scaffolds with CaO2 demonstrated sustained oxygen generation for 21 days in vitro and 28 days in vivo.	[82]
CaO2	Electrospinning	PLA	ADMSCs	-	Bone tissue engineering	Electrospun fibers supported adhesion, arrangement, and migration of MSCs and favorably influenced the alkaline phosphatase/osteocalcin expression and calcium deposition by MSCs	[85]
CaO2	Photocrosslinking	GelMA	SCAPs	-	Endodontic regeneration	0.5% CaO2 was sufficient to supply in situ oxygen for maintaining the embedded SCAP viability for 1 week. Moreover, hydrogels promoted SCAPs viability within root canal.	[87]
CaO2	Photocrosslinking	GelMA	CSPs	-	General application/cardiac tissue engineering	hydrogels supplied adequate oxygen, reduced metabolic stress of encapsulated cardiac progenitor cells, and increased the cell viability compared to plain GelMA hydrogels	[63].
CaO2	Emulsification/photocr osslinking	PCL	NIH/3T3 fibroblast s, L6 rat myoblasts, and primary cardiac fibroblasts	GelMA, PVA		In vitro studies suggest that GelMA hydrogels containing CaO2-loaded PCL microparticles were able to demonstrate oxygen release for approximately 5 weeks and supported the proliferation of three cell lines.	[73]
MnO2/CaO2	Phase separation/ionic crosslinking	PCL	endotheli al cells and smooth muscle cells	Glycerol, PVA, alginate	cardiac tissue engineering/skelet al muscle tissue engineering	In vivo study demonstrated that hydrogel scaffold was successful to provide sufficient oxygen amount to maintain high cell viability in aorta which provide prolonged survival times for transplanted tissues.	[133]
SPC/CaO2	Thermal crosslinking/curing agent	PCL, PVA		Gelatin, silicone, PVDC	Skin tissue engineering, wound healing	The in vivo full thickness wound model on pig skin noted that the four-layer film, promoted the expression of CD 31 which accelerated angiogenesis and woun healing.	[134]
SPC	Solvent casting	PLGA	-	-	Skin tissue engineering, wound healing	By day 3 and 7, the flap on in vivo models on mice, films with SPC have supported healing. However, after 7 days, the healing was found similar as control (without SPC).	[135]
SPC	Cryogenize d and sieved particles	-	C2C12 myoblast line	-	-	SPC preserved the contractility of skeletal muscle both in vitro and in vivo under hypoxia.	[136]
PFC	-	Fibrin hydrogel	MSCs	-	Bone tissue engineering, muscle tissue engineering	The hydrogel showed superior bone healing on in vivo mouse model when implanted on ectopic site compared to lumbar paravertebral muscle.	[137]
H2O2	Homogenized, emulsfied, sieved	-	Rat skeletal muscle cell	PLGA/Alginate, immobilize d catalase	Muscle tissue engineering	Increase in H2O2 percentage that is directly in contact of the cells cause toxicity to cells.	[138]

Abbreviations: PCL: polycaprolactone, nHA: nano-hydroxyapatite, SPC: Sodium percarbonate, PFC: perfluorocarbons, PDMS: polydimethylsiloxane, MnO₂: manganese dioxide, H₂O₂: hydrogen peroxide, SAP: self-assebling peptides, PLGA: poly(lactic-co-glycolic acid, GA: glutaraldehyde, MSCs: mesenchymal stem cells, PBMSCs : peripheral blood mononuclear-stem cells, ADMSCs: Adipose tissue-derived mesenchymal stem cells, SCAPs: stem cells from apical papilla, ADSC-EXO: adipose derived stem cell exosomes, CSPs: cardiac side population cells, PVA: poly(vinyl)alcohol, PVDC: polyvinylidene chloride, GelMA: gelatin methacryloyl.

Various studies have suggested that oxygen-generating materials could support stem cell proliferation and differentiation [83][84]. For example, hypoxic cell cultures corroborated that PLA scaffolds with CaO₂ supported adhesion, arrangement, and migration of MSCs and favorably influenced the alkaline phosphatase/osteocalcin expression and calcium deposition by MSCs [85]. Furthermore, CaO₂ nanoparticles loaded in PLGA microparticles were able to provide oxygen levels of 35–67.5 mmHg up to 14 days, improving the viability of the mesenchymal cells, and increasing alkaline phosphatase and osteocalcin expression [75]. The oxygen-generating electrospun scaffolds with CaO₂ nanoparticles in poly(glycerol sebacate) (PGS) and PCL demonstrated significant improvement in cell metabolic activity for primary bone marrow stem cells (BMSCs) [86]. Using these scaffolds, the cranial bone defects in rats healed significantly faster than those that did not generate oxygen. In another interesting study that used 3D printed oxygen generating scaffolds, transplanted constructs with preseeded BMSCs could reduce local cell apoptosis, enhance the survival of grafted cells, and support the osteogenic and angiogenic activity compared to the bare scaffolds without oxygen generation capacity [43]. In another recent study, an oxygen-generating composite scaffold that contained CaO₂-loaded gelatin microparticles were embedded in self-assembling peptides (SAP) and PLGA membranes. These constructs facilitated sustained oxygen generation for 21 days in vitro and 28 days in vivo [82]. The survival of the seeded peripheral blood-derived mesenchymal stem cells was significantly higher in the presence of oxygen-generating microspheres. The presence of the gelatin-CaO₂ microspheres significantly improved the repair of critical-sized cranial bone defects in another literature report (Fig. 5C–E). CaO₂-loaded oxygen-generating GelMA scaffolds were effectively used for endodontic regeneration [87]. A 0.5% CaO₂ concentration in GelMA was sufficient to provide the essential oxygen for maintaining the cell viability and improving the survival of stem cells from apical papilla (SCAPs) within the root canals of the teeth. Moreover, in a myocardial infarction model, a combination of stem cell derived exosomes and CaO_2-loaded scaffolds could improve cardiac function, reduce scar formation, and improve vascularization [88].

To recap, tissue engineering scaffolds combined with oxygen-generating materials have been shown to maintain physiologically relevant oxygen levels and support the survival of cells in engineered tissues after implantation. This aspect will help address the hypoxia-induced cell death which is a main challenge in ex vivo tissue engineering. In addition, the oxygen-generating scaffolds are anticipated to prevent hypoxia-induced cell death and support the clinical translation of cell-laden tissue engineered products.

6. Oxygen-Generating Scaffolds for In Situ Tissue Engineering

Ex vivo tissue engineering has significant limitations that cannot be solved by solely incorporating oxygen-generating materials. These include the health condition of the donors, the need for large numbers of donor cells to produce cell-seeded constructs, and the challenges posed by expensive and time-consuming in vitro cell culturing processes. Such challenges in ex vivo tissue engineering have led to the use of in situ tissue regeneration strategies, which utilizes the innate regenerative capacity of the body to reconstruct or repair the damaged tissues using cell-supporting scaffolds [89,90]. In situ tissue engineering has been shown to promote cellular colonization into implanted bioactive scaffolds and facilitate scaffold vascularization as well as tissue regeneration [91,92]. Cell-free scaffolds retain longer shelf life and are less demanding when being stored, transported, and handled, increasing the potential for clinical translation [93]. In such in situ tissue engineering approaches, a suitable microenvironment should be created in the scaffolds to ensure effective cell migration, survival, and incorporation of migrated cells into the implanted scaffolds [17,18]. As previously mentioned, oxygen is essential for cell growth, proliferation, and survival. It is critical to ensure the sustained delivery of oxygen in such tissue engineering approaches until the functional vascularization is achieved. Recent research on the utilization of oxygen-generating materials for in situ repair/generation of tissue defects has demonstrated that oxygen-generating scaffolds can supply sufficient oxygen to migrating and proliferating cells. In situ tissue engineering attempts with oxygen-generating materials can overcome the hypoxia-induced challenges revealing their remarkable and promising future in the clinic.

Recent studies have demonstrated that oxygen gradients can be generated to facilitate rapid migration of cells from the host tissue to the implanted scaffolds [94] (Fig. 6A). This behavior was demonstrated by subcutaneous injection of oxygen-generating and blank microparticles into mice. The oxygen-generating microparticles favored cell infiltration to the injection site (Fig. 6B). The development of fully functional vasculature and the establishment of oxygen and nutrient supply takes several weeks. Therefore, to ensure the survival of migrated cells on the scaffolds, it is of utmost importance to provide sufficient oxygenation. Previous studies indicated that oxygen-generating scaffolds could support attachment and migration of fibroblasts as well as promoting tissue healing minimizing necrosis [95]. A concise summary of major studies that used oxygen delivery approaches applicable for in situ tissue engineering are given in Table 3. In one example, a thin coating of CaO₂ in PCL was applied over 3D printed biphasic calcium phosphate (BCP) scaffolds. These constructs have provided sustained generation of oxygen at the implantation site and promoted bone ingrowth [96]. Similarly, in another study, oxygen-generating CaO₂-laden microparticles were loaded into 3D scaffolds to provide oxygen for up to 14 days, supporting cell migration, proliferation, and rapid repair of critical sized cranial defects in rats [42]. Upon subcutaneous implantation of CaO₂-loaded polyurethane cryogel scaffolds in mice, tissue necrosis was minimal in the CaO₂-loaded groups compared to the controls [41] (Fig. 6C). Furthermore, oxygen-generating cryogels displayed expression of PCNA, indicating viable and proliferating cells in the implanted scaffolds (Fig. 6D). Studies also showed that oxygen-generating and electrically conductive injectable hydrogels, based on gelatin-graft-polypyrrole and periodate-oxidized pectin, are promising scaffolds for in situ tissue engineering [97].

Fig. 6.

Oxygen-generating materials for in situ tissue engineering. A) A scheme showing the effect of oxygen-generating materials on cell survival, especially in the deeper regions of the scaffold upon implantation. B(i) Photographs and B(ii) Hematoxylin and eosin (H&E) staining images of mouse skin tissue after subcutaneous injection of test microparticles. B(iii) corresponding quantification of infiltrating cells. Asterisks indicate the implanted microparticles or the areas originally occupied by microparticles. C) Quantification of percent necrosis after day 3 and day 9 of implantation of oxygen-generating and blank scaffolds. D) H&E staining of the skin flaps harvested at [D(i)] day 3 and [D(ii)] day 9 showed increased necrosis in the PU group, whereas necrosis was minimal and skin architecture was maintained in the PUAO and PUAO-CPO groups. Picrosirius red staining showing the collagen architecture of harvested skin at [D(iii)] day 3 and [D(iv)] day 9 of implantation. Fig. B is reproduced from Ref. [94] with Creative Commons Attribution License (CC BY 4.0). Fig. C and D are reproduced from Ref. [41] with the permission American Chemical Society.

Table 3.

Oxygen-generating scaffolds for in situ tissue engineering.

Oxygen-generating material	Fabrication method	Scaffold material	Other components	Application	Outcome	Ref.
CaO2	Freeze-drying	Collagen/Chitosan	Ciprofloxacin	Skin tissue engineering	Suitable cell attachment and migration for fibroblasts (in vitro). Better wound healing and less necrosis (in vivo).	[95]
CaO2	Direct-write assembly (robocasting)	Beta-tricalcium phosphate (ß-TCP) / Hydroxyapatite (HA)	PCL	Bone tissue engineering	Increased cell viability and promoted bone ingrowth.	[96]
CaO2	UV crosslinking	GelMA	PCL	Bone tissue engineering	Increased cell viability, proliferation, and cytocompatibility.	[97]
CaO2	Cryogelation	Antioxidant polyurethane (PUAO)	-	Wound healing and cardiac tissue engineering	Increased metabolic activity and cell viability. (in vitro) Delayed onset of necrosis. (in vivo)	[41]
H2O2	Thermal gelation	Gelatin-graft-polypyrrole/Periodateoxidized pectin	PLA	Bone tissue engineering	Increased cell viability and attachment.	[97]
CaO2	Electrospinning	Antioxidant polyurethane (PUAO)/collagen	exosomes	Cardiac tissue engineering	Promoted angiogenesis and heart regeneration.	[88]
PVP/H2O2 complex	Thermal Gelation	PNIPAAm	PLGA	Cardiac tissue engineering	Increased cell survival and proliferation, greater angiogenesis, reduced inflammation, and decreased cardiac fibrosis	[139]
CaO2	Ionic crosslinking	Alginate	-	Skin regeneration	Improved tissue infiltration, wound closing, and wound restoration	[140]
CaO2	Cryogelation	Antioxidant polyurethane (PUAO)	-	Skin tissue engineering	Prevented necrosis	[41]
CaO2	Casting	Polyacrylamide/Sodium alginate	Vitamin C	Bone Tissue engineering	Promoted Zhangbone healing	[66]
CaO2	Casting	Silk fibroin/keratin	Gelatin	Urinary Tract Tissue Engineering	Improved organized muscle bundles and the epithelial layer	[141]
CaO2	Electrospinning	PCL/Gelatin	Polydopamine	Wound healing	Promoted wound healing	[142]
SPC	Electrospinning	PCL	-	Wound healing	More vascularization and wound healing potential	[143]

Abbreviations: CaO₂: Calcium peroxide: SPC: Sodium percarbonate: H₂O₂: hydrogen peroxide, PLGA: poly(Lactide-co-Glycolide),, PCL: polycaprolactone, PLA: polylactic acid, PVP: polyvinylpyrrolidone, PNIPAm: Poly(N-isopropylacrylamide)

7. Oxygen-Generating Bioinks for 3D Printed Tissue Constructs

3D bioprinting is an additive manufacturing method that allows for fabrication of complex 3D tissue structures in the laboratory to repair and reconstruct damaged tissues [98]. 3D bioprinting strategies mainly involve three steps; the generation of a CAD model of the tissue to be developed, formulation of a suitable bioink, and the printing of the construct using a 3D bioprinter [2]. Different types of biomaterials, cells, and growth factors/chemical cues have been used in bioink formulations based on the specific requirements [99]. Bioprinting involves layer-by-layer deposition of the bioink according to the CAD design [21,100]. Various crosslinking strategies including ultraviolet based and visible light-based ones are used in 3D bioprinting to generate 3D constructs [101].

Providing oxygen to the embedded cells in the bioprinted constructs is essential to maintain high cell viability in the printed tissue structures. Passive oxygen transport is only suitable at short distances, limiting effective oxygen transport in large, engineered tissues [102]. Larger constructs require vascularization to provide adequate oxygen because of the large distances between the cells and the nearest capillaries. A recent study indicated that oxygen perfusion in 3D bioprinted scaffolds with microchannel networks maintained high oxygen levels in cell-laden constructs and improved cell viability [103].

However, developing vascularized constructs using 3D bioprinting is challenging. Despite the recent progress, creating highly complex vasculature in tissue constructs is a demanding task using the currently available technologies [104]. Due to the limited resolution of bioprinting, fabrication of blood capillaries in the bioprinted constructs can be challenging [105,106]. The rate of functional blood vessel formation upon implantation is not rapid enough to provide sufficient oxygen to the tissue [107]. In the biofabricated scaffolds that contain cell-laden bioinks, apoptosis can occur for the cells that are located deep inside the construct due to hypoxia [106,107]. To compensate the time required to form vasculature, oxygen-generating materials can be utilized to provide oxygen to the cells [40]. Moreover, these materials can support the development of vasculatures by providing sufficient oxygen to the cells required for vascular sprouting [108,109].

The scalability of 3D bioprinted structures is another obstacle to address. Attempts to scale up bioprinted structures have yielded in unsatisfactory outcomes when large tissue constructs were used [110]. One of the major reasons for the failure of large scale bioprinted constructs is the limited oxygen diffusion into the bulk of the construct. Oxygen-generating materials embedded in printed constructs can solve this issue by providing sustained oxygen supply to the cells.

Several recent attempts utilized solid peroxides like CaO₂ as oxygen-generating materials in bioprinted constructs (Fig. 7A). Table 4 summarizes the recent approaches that used various oxygen generating materials in bioprinting. In a recent study, 0.5% CaO₂ was combined with GelMA bioink to develop oxygenated bioprinted constructs [34]. Fig. 7B indicates that 0.5% of CaO₂ had the highest shape fidelity and a thinner mesh structure. Results showed that cardiomyocyte-laden GelMA with CaO₂ increased cell survival and function under hypoxia, as seen in Fig. 7C, and 7D. In another study, a GelMA hydrogel and CaO₂ based bioink formulation was used for developing bioprinted skeletal muscle tissues [111]. Results revealed that the metabolic activity, proliferation, and viability of the mouse derived C2C12 myoblast cells were higher when the oxygen-generating material was present. The data revealed that 0.5 mg/mL CaO₂ was the optimum concentration for maintaining adequate cell viability and metabolic activity in the printed constructs.

Fig. 7.

Examples of oxygen-generating materials in bioprinting. A) Schematic representation of an oxygen-generating biomaterial and bioink composition. B) Microscopic images that show the variation in pore morphology, pore size, and strand size with different concentrations of calcium peroxide (CPO) in the bioprinted scaffolds. i) 0% ii) 0.1% iii) 0.5% iv) 1%. C) Quantification results for cell viability (*p < 0.05; **p < 0.01). D) Fluorescence images for live/dead staining of the 3D bioprinted cardiomyocyte-laden GelMA constructs with 0% and 0.5% CPO concentration under hypoxic and normoxic conditions at different time points (Scale bar is 200 μm). E) Histological analysis of 3D printed oxygen-releasing structures which have 0.3 mg/mL CPO concentration at day 3 and day 6. Cells that undergo mitosis are shown by black arrows. Figure A is created with BioRender.com. Fig. B, C, D are reproduced from Ref. [34] with the permission of Wiley. Fig. E is reproduced from Ref. [84] with the permission of IOP Publishing.

Table 4.

Oxygen-generating bioinks for 3D bioprinted tissue constructs.

Oxygen-generating material	Biopolymer/hydrogel	Bioprinting method	Cell type	Application	Outcome	Ref.
CaO2	GelMA	Extrusion	NIH 3T3 Fibroblasts and Cardiomyocytes	Cardiac tissue engineering	Metabolic activity and viability of cells improved	[34]
CaO2	GelMA - Alginate	Extrusion	C2C12 Myoblasts	Skeletal muscle tissue engineering	Metabolic activity of cells improved	[111]
CaO2	Alginate	Extrusion	ADSCs	General	Cell viability and proliferation improved. In vitro Hypoxia-induced apoptosis was reduced.	[84]
Chlamydomonas reinhardtii (a type of green algae)	GelMA	Extrusion	HEPG2, C2C12 mouse myoblasts, and HUVECs	Liver tissue engineering	Hepatic tissue constructs-maintained viability and function	[114].
Chlorella pyrenoidosa (a type of microalgae)	Alginate - GelMA	Extrusion (in situ)	Human skin fibroblast (HSFs)	Wound healing	Increased chronic wound closure in vivo	[144]
SPO and CaO2	PCL and fibrinogen-based hydrogel	Extrusion	C2C12 cells	Skeletal muscle tissue engineering	Increased cell viability and further differentiation to elongated myotubes	[145]
H2O2	GelMA	Extrusion	C2C12 cells	Tissue engineering	Increased cellular survival and morphology	[68]
CaO2	Agarose-Alginate	Extrusion	E. coli	Bacterial research	Expression of fluorescent proteins deep within the gel	[146]
CaO2	Non-mulberry silk, PEGDMA-GelMA (with single-walled CNTs)	Extrusion	HUVECs	Cardiac tissue engineering	Implantable 3D bioprinted cardiac construct for an infarcted region in the heart	[147]

Abbreviations: CaO₂: calcium peroxide, SPC: sodium percarbonate, H₂O₂: hydrogen peroxide, ADMSCs: adipose tissue-derived stem cells, PCL: polycaprolactone, GelMA: gelatin-methacrylamide, PEGDMA: polyethylene glycol dimethacrylate, HUVECs: human umbilical vein endothelial cells.

Stem cells are promising cell sources for bioprinting due to their self-renewal, proliferation, and differentiation properties [112,113]. Adipose-derived stem cells have been used in an alginate based bioink that contained CaO₂ to ensure the survival of cells in the construct [84]. The CaO₂ particles provided a sustained release of oxygen and increased the viability of the cells in the printed constructs. Oxygen-generating 3D bioprinted constructs maintained a suitable microenvironment for cell proliferation and diminished hypoxia-induced apoptosis in vitro (Fig. 7E).

In addition to the approaches that use inorganic peroxides, recent studies focused on developing bioprinted constructs loaded with other types of oxygen-generating components. For example, researchers have demonstrated that Chlamydomonas reinhardtii, a type of green algae, could be used as a source of oxygen [114]. They co-cultured photosynthetic C. reinhardtii and mammalian cells to print perfusable vascularized tissue constructs. The C. reinhardtii cells were embedded into GelMA and used as oxygen source to the tissue that contained liver-derived cells (HepG2). Results demonstrated that the incorporation of C. reinhardtii in the bioprinted structures improved the viability and functionality of HepG2 cells.

The selection of the appropriate bioprinting technique and suitable types of cells, biomaterials, and chemical factors are crucial for the successful development of engineered tissues [106]. The printability of the material is the primary parameter that determines the mechanical, chemical, and biological properties. The addition of oxygen-generating materials into bioinks could change the printability properties because of the alteration of viscosity [34]. Furthermore, the size and architecture of the printed construct are important parameters controlling the release kinetics of oxygen. Thus, it is crucial to explore new 3D printing approaches, new bioink formulations, and controlled oxygen delivery approaches to develop clinically relevant constructs.

8. Existing Challenges and Prospects of Oxygen-Generating Materials in Tissue Engineering

As in any new field, the success of oxygen-generating materials in tissue engineering depends on addressing various challenges. An adverse effect can be expected when a foreign material is introduced to the body particularly at higher concentrations. Current advances in biomedical research focus on minimizing such adverse effects while enhancing effectiveness through innovative design approaches. The major challenge in the oxygen-generating biomaterials field is the toxicity of the materials used for oxygen delivery, by-products and intermediates produced during oxygen generation, and the local adverse effects due to the uncontrolled oxygen supply.

Most of the oxygen-generating materials are selected based on the previously reported literature results in other fields such as in the environmental sciences, soil remediation, aquatic life management, and catalysis. Directly adopting the concept of oxygen generation from such fields and implementing it in therapeutics can cause an overlook of the potential risks and microlevel adverse effects. CaO₂ was deemed relatively safe for oxygen generation in environmental remediation. However, when it is directly implanted in the body, the effects in the human body are significantly different. For instance, calcium and magnesium peroxides may affect the skin and eyes upon direct contact [115][116]. The synthesis and handling of these peroxides should be carefully performed as they may cause irritation to the lungs, nose, and throat if inhaled or ingested [116].

When using as a pharmaceutical agent, detailed studies are required to identify the therapeutic window of each oxygen-generating material to fully utilize their potential without detrimental effects. Moreover, detailed genotoxicity, teratogenicity, and systemic toxicity studies are also required before the clinical application of oxygen-generating materials can take place. Additionally, absorption, distribution, metabolism and excretion (ADME) studies of each oxygen-generating agent must be performed to identify a safe amount that can be used in treatment regimens, to determine the required time between each use, and the chances of biomagnification in the body. In addition, the release of metal ion from the oxygen-generating agent should be minimal to avoid metal ion associated harmful effects. For instance, excess calcium ions and magnesium ions in the blood may cause pathophysiological conditions such as hypercalcemia (total serum calcium level is above 10.4 mg per dL) and hypermagnesemia (total serum magnesium level is above 2.6 mg per dL), respectively [117,118]. Thus, robust strategies should be developed and adopted to control the release of such ions to minimize the adverse manifestations in the body.

In addition to the direct effects of oxygen-generating agents in the body, the intermediates of oxygen formation reaction can also result in adverse effects. During the hydrolysis of solid peroxides to generate oxygen, H₂O₂ is produced as an intermediate [119]. Apart from various important roles of H₂O₂ in wound healing, wound protection, angiogenesis and tissue regeneration, it can induce necrotic cell death [120]. Thus, limiting the generation of H₂O₂ and increasing the conversion speed and efficiency of H₂O₂ into oxygen are required to minimize the detrimental effects [121]. Superoxide toxicity is generally attributed to its ability to reduce metal ions. Consequently, reoxidation of the metal by H₂O₂ yields deleterious oxidizing species. Reactive oxygen species (ROS) serve as cell signaling molecules for biologic processes [58]. However, the generation of ROS can also damage some of the cellular organelles and processes, ultimately disrupting normal physiology [122]. As discussed in this review, antioxidant enzymes such as catalase can be utilized to minimize the unwanted effects of radicals. Recent studies demonstrate that incorporating catalase enzyme into 3D scaffolds favors the conversion of H₂O₂ into oxygen minimizing the adverse effects [37,97]. Moreover, the formation of H₂O₂ and other ROS species from oxygen-producing substances has a greater impact in vitro than the effect in vivo. Due to a limited number of cells being utilized, the ability of such cells to produce catalase or other enzymes is limited, resulting in a greater impact in in vitro than in vivo. The use of limited number of cells in vitro would not be as efficient as using them in a heterogeneous population in vivo. It is evident that cells containing peroxisomes, such as liver and kidney cells, are able to produce a higher amount of such enzymes than other cells [123]. Thus, a potential mismatch between the in vitro and in vivo results in toxicity and oxygen release data might occur. Moreover, free radicals can also be generated from H₂O₂ in the presence of metal ions like copper or iron in the body. Similarly, free radicals can form after ischemia or reduced blood flow to an organ by the mitochondria. A reperfusion mimetic condition that brings more oxygen takes place when the ischemic tissue is supplied with an engineered construct with oxygen generation potential. Oxygen might react with the pre-existing free radicals to form more radicals possibly causing cell damage. In addition to the currently employed approaches that use catalase, other agents such as antioxidants (vitamin A, vitamin C, and vitamin E) can donate electrons to neutralize the free radicals and protect cells. Glutathione, a biological molecule produced in the body, acts as an antioxidant and neutralizes H₂O₂. Antioxidants can possibly be incorporated into oxygen-generating scaffolds to prevent the adverse effects and cell death in vitro due to the ROS, peroxide, and other radicals generated. Incorporating catalase in polymeric scaffolds is still a challenge mainly because the enzymatic activity can be compromised upon dissolving catalase in organic solvents. Moreover, harsh processing parameters employed during scaffold development can also limit incorporation of catalase into scaffolds. For instance, 3D printing utilizing high temperature for processing (fused deposition modeling, FDM) may not be able to incorporate thermally labile molecules. To overcome such challenges, strategies for incorporating catalase in nanostructured materials or relatively stable polymeric nanocarriers can be a promising approach.

Despite the promising results of in vitro studies that use oxygen-generating materials in bioprinting, these approaches might be limited due to the nature of the bioprinting processes and the complexity of tissue vasculature. Since bioinks are based on aqueous solutions, there is a higher chance of inducing oxygen generation during the bioink preparation steps. To prevent the pre-mature release of oxygen, novel encapsulation methods should be developed. Approaches like combining other fabrication techniques like electrospinning can be a promising strategy to facilitate controlled release [124]. In addition, bioprinting utilizes technical knowledge from various other fields, and advancements in these areas can also accelerate optimization of bioprinting strategies. For example, induced pluripotent stem cells (iPSC) have been used in bioprinting studies [125]. However, spatiotemporal effects of oxygen-generating materials in stem cell fate within bioprinted constructs need to be further studied.

Excess formation of oxygen in local tissues can result in toxic manifestations posing a challenge in development of oxygen-generating materials [126]. Studies indicated that among patients who survived after a cardiac arrest with hyperoxygenation, a majority of them was associated with mortality after a second cardiac arrest [127]. A clinical study showed that even brief exposure to high oxygen concentration can induce oxidative stress in leukocytes and platelets [128]. Although such reports should be supported by further research and analysis, the risk of hyperoxygenation and adverse effects should be consistently expected. Designing biomaterial scaffolds loaded with controlled amounts of oxygen-generating materials with tightly modulated release kinetics could solve this challenge. Recent research also indicated that with the help of advanced material fabrication approaches utilizing stimuli-responsive carriers, it is possible to deliver oxygen to the tissues in a precisely controlled manner. This outcome could possibly be achieved by varying the temperature, pH, ratio of peroxide to water, volume of catalyst, and the type of catalyst because these factors influence the rate of oxygen formation from peroxide compounds. The release of oxygen can be controlled externally in the implanted tissue engineered constructs by applying external stimuli such as light, magnetic forces, and electrical stimulation as well. In addition to the ongoing in vitro and in vivo studies using oxygen-generating tissue engineered constructs, future research should focus on translation into clinic and clinically relevant products.

To summarize, tissue engineering products can potentially be translated into clinic with further research to overcome the current challenges of oxygen-generating materials.

9. Conclusions

Successful development of tissue-engineered products containing oxygen-generating materials is anticipated to significantly improve the regeneration and repair of defective tissues by preventing the otherwise inevitable hypoxia-induced cell death. Recent approaches have been developed to mitigate the adverse effects of peroxide radicals generated during the oxygen generation process. Furthermore, strategies have been developed to provide a sustained release of oxygen throughout the regeneration process are highly encouraging. The applications of oxygen-generating materials span from ex vivo and in situ tissue engineering to 3D bioprinting. In ex vivo tissue engineering, oxygen-generating materials help support cells during the in vitro culture of cells in the scaffolds as well as in their in vivo implantation. Similarly, oxygen-generating materials help promote cell migration and provide a suitable microenvironment for cell survival upon implantation in in situ tissue engineering. The incorporation of oxygen-generating materials in bioinks can also satisfy the need for oxygen in 3D bioprinted large structures. Although some challenges including the risks of cytotoxicity associated with some of the oxygen-generating materials still exist, there are ongoing innovative approaches to minimize such adverse effects. Oxygen-generating biomaterials are anticipated to enhance the bench-to-bed translation of tissue-engineered products over time.