Breathing life into engineered tissues using oxygen-releasing biomaterials

Engineering three-dimensional (3D) tissues in clinically relevant sizes have demonstrated to be an effective solution to bridge the gap between organ demand and the dearth of compatible organ donors. A major challenge to the clinical translation of tissue-engineered constructs is the lack of vasculature to support an adequate supply of oxygen and nutrients post-implantation. Previous efforts to improve the vascularization of engineered tissues have not been commensurate to meeting the oxygen demands of implanted constructs during the process of homogeneous integration with the host. Maintaining cell viability and metabolic activity during this period is imperative to the survival and functionality of the engineered tissues. As a corollary, there has been a shift in the scientific impetus beyond improving vascularization. Strategies to engineer biomaterials that encapsulate cells and provide the sustained release of oxygen over time are now being explored. This review summarizes different types of oxygen-releasing biomaterials, strategies for their fabrication, and approaches to meet the oxygen requirements in various tissue engineering applications, including cardiac, skin, bone, cartilage, pancreas, and muscle regeneration. Biomaterials that surround artificial tissue and slowly release oxygen can improve survival rates during wound healing and organ regeneration trials. Although researchers have successfully constructed many kinds of tissues using 3D scaffolds, supplying the cells with adequate oxygen and nutrients post-implantation remains challenging. Gulden Camci-Unal from the University of Massachusetts Lowell in the USA and colleagues review how microscale oxygen sources such as calcium peroxide crystals are being combined with flow-regulating polymers to keep engineered tissues healthy until vascular systems are established. Coatings of peroxides and polymers applied to 3D scaffolds, for example, strengthen bone cultures by releasing a steady supply of oxygen to permeate throughout the tissue for extended times. Byproducts from oxygen-releasing materials can also be used to inhibit biofilm growth on implants, lessening chances of infection. The inability to administer oxygen in a controlled and sustained manner into thick artificial tissues has attracted a growing interest towards the design and development of new functional biomaterials. Without a sufficient oxygen supply, tissues suffer from the effects of apoptosis and necrosis. Incorporation of oxygen-releasing materials into scaffolds can help address this challenge. This paper provides an overview of the recent developments and technological advances in engineering oxygen-releasing biomaterials to improve the viability and function of cells and prevent hypoxic tissue death. Recent advances in different types of oxygen-releasing materials, mechanisms of oxygen generation, and their applications are discussed.


Introduction
For many years, the reviews in this field have summarized the developments and applications of different oxygen-releasing biomaterials. Previously published examples have discussed applications of oxygen-releasing biomaterials for a limited number of tissues. However, major developments in the field have occurred more recently to expand on the current knowledge of this research area. This review provides a comprehensive overview of oxygen-releasing biomaterials, including the most recent developments and literature examples of new and innovative uses of oxygen-releasing constructs. Extensive tissue engineering applications in vitro and in vivo are elaborately discussed, including their different synthesis methods, oxygen release mechanisms and release kinetics, critical insights, and possible future applications.
This section provides an overview of the current challenges in the tissue engineering of 3D constructs and discusses the current approaches and biomaterials used to address these limitations. Different types of oxygenreleasing biomaterials, their synthesis and fabrication methods, and release kinetics and mechanisms are examined.

Tissue engineering: current need and limitations
Organ damage and failure are the leading causes of death globally 1 . According to the United States Department of Health and Human Services, more than 116,000 people in the United States were on waiting lists for organ transplants in August 2017. Since the availability of matching donors does not always meet the demand of organ transplants, there are approximately 20 deaths each day because of delayed organ transplantations 1,2 . In 2015, 119, 363 people in the United States were in need of an organ donation, which is double the statistics recorded in 1999 1,3 . Tissue engineering approaches can address the unmet demand for the replacement, repair, and regeneration of organs and tissues [4][5][6][7] . There has been remarkable success in the clinical translation of 3D tissueengineered constructs for repairing organs, such as the trachea 8,9 and lungs; 10 however, multiple challenges remain unaddressed. These challenges include (i) the preservation of high cell viability postimplantation, (ii) the unwanted immune response, (iii) the obstruction of cell signaling pathways, (iv) insufficient cell proliferation, and (v) low metabolic activity within the implanted scaffolds. In addition, currently available approaches do not yield properly functioning engineered tissues because they cannot provide sustained amounts of oxygen 11 .
Ideally, functional tissue constructs must deliver a supply of oxygen, nutrients, and growth factors to the encapsulated cells in 3D scaffolds. These components ensure that the cells survive, proliferate, differentiate and remain metabolically active. In the human body, the primary source of nutrients for tissues is the host blood supply through the vascular system 12 . Generally, complete vascularization requires gradual development over time postimplantation 12,13 . The vascular system also includes networking for the removal of cellular wastes from tissues 14 . Therefore, engineered scaffolds that can offer these chemical and structural properties tend to support high tissue viability. The use of oxygen-releasing biomaterials may overcome the oxygen diffusion limitations in organ-sized constructs 15,16 . Until such optimum vascularization is achieved in vivo, oxygen-releasing materials may provide an adequate oxygen supply to support the viability and metabolic function of these tissue constructs. Addressing these considerations could be a promising and an impactful solution to the clinical translation of tissue engineered constructs 15,17 . In conjunction with modern vascularization strategies, the removal of cellular wastes and the delivery of growth factors is also accomplished, leading to a highly biomimetic scaffold 18,19 .
Engineered tissues that are not within the diffusion and mass transport range for effective vascular integration do not usually survive 11 . There is a period of 4-6 weeks for vasculature to reach approximately 83% patency within a damaged tissue 20 . During this time, the capillaries and blood vessels of the host integrate and grow into the implanted engineered tissue grafts. The viability of engineered tissues is highly dependent on the amount of oxygen available during integration 21 . However, the solutions to support high oxygen demands and maintain high viability during this critical developmental stage are still in their infancy. Therefore, new biomaterial innovations must address the insufficient oxygen supply in current tissue engineering methodologies.
Tissues with high oxygen demand: why the clinical translation of tissue engineering approaches fail?
Without a sufficient supply of oxygen, tissues undergo necrosis and programmed cell death [22][23][24][25] . The tendency to undergo necrotic death is especially evident for cells and tissues with high metabolic activity and oxygen demand. Maintaining the viability of such tissues (e.g., cardiac, pancreas, muscle) is a challenge that has prolonged the clinical translation of engineered tissues for the treatment of diseased or damaged organs 11 . Perfusion in conjunction with convective diffusion mechanisms is a method to improve oxygen availability for engineered tissues but has limited success 26 . The extent and speed of optimum vascularization in tissue-engineered grafts depends on multiple factors: scaffold design, porosity, interconnectivity, the influence of growth factors, and paracrine signaling, which facilitates tissue development [27][28][29][30] . These factors motivate the development of functional materials that mimic the in vivo microenvironments and facilitate the functional development and integration of engineered tissues 31 . As a result, innovative approaches to fabricate thick and vascularized constructs are currently studied for the clinical translation of engineered tissues.

Current strategies for the survival of tissue-engineered constructs with high oxygen demand
For tissues with greater oxygen demands, various approaches have been utilized and shown to either increase the oxygen content or functionality of tissue constructs. Supplementing the cell media with oxygen carriers has enabled the increase in oxygen content in tissue engineered constructs 32 . Other technologies, such as perfusion bioreactors, can culture cells in welloxygenated environments. To improve vasculature and tissue viability, cell-laden samples have been treated with growth factors, such as vascular endothelial growth factor (VEGF). Although promoting vasculature is important, the effectiveness of this strategy is limited by the inhomogeneous incorporation of growth factors in 3D 33 . Microfluidic techniques for the fabrication of vascular networks within 3D scaffolds also have allowed the guided development of blood vessels. Despite the in vitro success of microfluidic approaches, the resolution of microfabrication techniques was not optimal 34,35 . Other studies have attempted to incorporate angiogenic cells into scaffolding matrices to promote rapid neovascularization [36][37][38][39] . The use of terminally differentiated endothelial or human umbilical vein endothelial cells (HUVECs) has been shown to have poor vasculature development both in vitro and in vivo 40 . To facilitate blood vessel formation, some approaches have also used porous interconnected scaffolds that release growth factors [41][42][43][44] . Until uniform blood vessel networks are formed, the integration with the host vasculature is incomplete and oxygen availability is restricted. Given the extensive timeline for vasculature integration, oxygen demands must be provided through external means. Therefore, biomaterial scaffolds must both consider the necessary networking and, importantly, the oxygen supply required. Recent progress in this research has recently shifted to biomaterials that can provide this oxygen supply over extended periods of time.

Approaches to engineer oxygen-releasing biomaterials
The transition from refining vascularization techniques to engineering oxygen-releasing biomaterials has seen significant progress over the past decade 11 . Oxygenreleasing biomaterials can be fabricated through the incorporation of solid peroxides, liquid peroxides, and fluorinated compounds into the scaffold polymer. These scaffolds can serve as 2D tissue culture substrates or be used in the 3D encapsulation of cells. Inorganic solid peroxides, such as calcium and sodium peroxides, liquid peroxides, and perfluorocarbons, have been used to supply oxygen to cells [45][46][47] . The most commonly used techniques to incorporate compounds for oxygen delivery are adsorption onto fiber-based scaffolds, 3D encapsulation in polymers, and incorporation into hydrogel networks 16,48,49 . Micro-and nanoparticles of solid peroxides (e.g., calcium, magnesium, and sodium peroxide) interact with water and undergo hydrolytic degradation to release oxygen 50 . The introduction of these particles has shown significant in vitro and in vivo successes. This review summarizes the different types of materials used to obtain the controlled release of oxygen in engineered tissues to prevent cell death due to hypoxic stress. Table 1 outlines the different types of oxygen-releasing compounds that are used in tissue engineering applications and their corresponding oxygen-release mechanisms.
Oxygen-releasing biomaterials: advantages, synthesis, composition, and characterization The need for oxygen-releasing materials has led to the development of new substrates that release oxygen in a sustained and controlled manner. Controlling the oxygen-release kinetics can significantly influence cell differentiation, viability, and proliferation. Solid inorganic peroxides, liquid peroxides, and fluorinated compounds can be used for sustained oxygen release and maintain high cell viability in the tissue construct 21 . The strategy of doping scaffolding polymers with oxygen-releasing peroxides and fluorinated compounds has been increasingly used to enhance the viability of tissue constructs 11 . In this section, we discuss the different types of solid and liquid oxygen-releasing reagents, approaches to incorporate these reagents into biomaterials, and their oxygengenerating mechanisms.

Solid inorganic peroxides
Solid inorganic peroxides have been successfully used as oxygen-releasing materials in vitro and in vivo. The most commonly used solid peroxides are sodium, calcium, and magnesium peroxides 16,50,51 . The main mechanism behind the oxygen release in these inorganic compounds is hydrolysis. When nano/microparticles of calcium, sodium, and magnesium peroxide interact with water, the particles undergo hydrolytic decomposition, as shown in the following equations [52][53][54] .
(Calcium peroxide) CaO 2 (s) Solid inorganic peroxides have been used to support cell growth, survival, tissue regeneration, and bioremediation. Among these materials, calcium peroxide has been shown to produce the highest amount of oxygen in a variety of applications 55 . The amount of oxygen released by inorganic solid peroxides depends on their solubility coefficients. The equilibrium coefficients of calcium and Muscle regeneration 34 , intervertebral disc regeneration 106 . magnesium peroxide are 9.8 × 10 −11 and 1.8×10 −11 , respectively 55 . Calcium and magnesium peroxide have low solubility coefficients (1.65 g/L at 20°C and 0.86 g/L at 18°C, respectively) 55 . Based on these values, calcium peroxide has a higher oxygen-generation potential than magnesium peroxide [55][56][57][58] . Magnesium peroxide produces the lowest amount of oxygen among solid peroxides 55 with a content purity of 15-20% by weight 11 . This effect is mainly due to the low solubility of magnesium peroxide in water, which reduces the rate of oxygen formation. While these approaches facilitate oxygen generation, burst release may occur at uncontrolled rates. This release behavior is especially found in hydrogels that have solid peroxides encapsulated without any hydrophobic barriers 11 . Table 2 summarizes the solubility coefficients of different solid inorganic peroxides. For more controlled rates of oxygen release, hydrophobic substrates can be utilized for the encapsulation of oxygen-releasing particles. Hydrophobic materials, such as polydimethylsiloxane (PDMS) and poly(D, L-lactidecoglycolide) (PLGA) act as barriers to reduce the rate of hydrolysis of the oxygen-releasing particles and help establish a more controlled rate of oxygen release 59,60 . Conversely, if the oxygen-releasing particles are surrounded by hydrophilic polymers, then the material allows for the rapid diffusion of water molecules into the 3D matrix, which can dramatically increase the rate of oxygen generation 11,61 . The rate at which oxygen is released from the solid peroxides is also dependent on other environmental factors, such as the pH, temperature, and peroxide-to-water ratio.
Solid inorganic peroxides readily undergo hydrolytic decomposition upon reaction with the water content in the biomaterial. The solubility of peroxide is a factor that can affect the overall oxygen-release kinetics and quantities in implanted constructs 16 . However, the release kinetics of solid inorganic peroxides are controlled by changing the temperature, pH, purity, and solubility of the compound. Consequently, the oxygen-release kinetics can be tailored to a wide spectrum of tissue engineering applications. The caveat is that some solid inorganic peroxides, such as calcium peroxide, can produce reactive oxygen species during hydrolytic decomposition. These byproducts are potentially harmful to the cells encapsulated in the scaffolding material. As an improvement, the addition of enzymes can assist in the decomposition reaction. Specifically, the catalase enzyme can decompose the intermediate hydrogen peroxide into water and oxygen. This intermediate step eliminates any potential cytotoxic reactive oxygen species 16 .

Liquid peroxides
In addition to solid oxygen-releasing reagents, liquid peroxides have been used in various biomedical applications. The higher solubility of liquid peroxides in water permits rapid oxygen release. Hydrogen peroxide is a liquid compound that can release oxygen upon decomposition into oxygen and water. In the human body, the enzyme catalase, present in the liver and blood, breaks down hydrogen peroxide into water and oxygen 62 . Although the actual role of catalase in the oxygen-release process is unclear, decomposition is suggested to take place through the following reactions: 63 In the absence of catalase, cytotoxic side products of this reaction, such as free hydroxyl radicals, are often generated and detrimental to the cells 64 . Liquid peroxides have also been used to control the efficiency of oxygen release in cell encapsulation approaches. For example, hydrogen peroxide was microencapsulated within PLGA and coated with a secondary layer made of an alginate hydrogel 64 . Then, catalase was chemically modified into the alginate backbone. The hydrogen peroxide is released from the PLGA and diffuses into the alginate layer. The alginate spheres containing catalase assisted in the decomposition of hydrogen peroxide into oxygen and water without generating harmful radicals that may have otherwise influenced cell viability 64 . Other strategies that used hydrogen peroxide as an oxygen source have demonstrated oxygen release over a two-week period in which successful cell differentiation was induced 11 . Figure 1 represents the chemical structures of different oxygen-releasing compounds. The commonly used solid peroxides, liquid peroxides, and fluorinated oxygenreleasing compounds are shown. The listed solid and liquid peroxides undergo rapid hydrolytic decomposition and release oxygen as a byproduct. The fluorinated compounds have a high oxygen-carrying capacity and release oxygen by diffusion mechanisms.
The rate of oxygen release is dependent on the efficiency of the liquid peroxide used in the application. In the absence of enzymes, liquid peroxides, such as hydrogen peroxide, can generate cytotoxic side products. To mitigate this issue, enzymes can assist in the conversion of hydrogen peroxide to water and oxygen. Specifically, catalase is a useful enzyme to achieve this result and contains iron and organic functional groups that enhance the oxygen-release capabilities. Moreover, catalase has a high turnover efficiency, which can improve the potency of oxygen-releasing compounds 11 .

Fluorinated compounds
Fluorinated compounds have functioned as oxygen carriers in numerous applications in cosmetics, drug delivery, protein stabilization, and organ preservation where the oxygen content is critical for the effectiveness of the product or the process 65 . In the last decade, the use of fluorinated compounds in tissue culture applications and their effectiveness in supporting the high oxygen demand in tissues in vitro and in vivo have become widely used in biomedical research 66 . In particular, perfluorodecalin, a fluorocarbon, in water emulsions have acted as an oxygen carrier and prevented ischemia after surgical procedures 65 . The surface chemistry of fluorinated surfactants, such as perfluorooctanesulfonic acid (PFOS) and perfluorooctanoic acid (PFOA), have demonstrated greater surface activity than that of their fluorocarbon analogs, such as perfluoroalkylated amine oxide 65 . The molecular structures of these surfactants, such as the fluorocarbon/hydrocarbon amphiphile FnHm (i.e., C 6 F 13 C 10 H 21 ), enable these compounds to have a self-assembling nature and form highly stable emulsions and gels. This feature allows control over the diffusion and oxygen-release kinetics, a desired property of oxygenreleasing materials. Some fluorinated surfactants, such as perfuoromethyl-cyclohexylpiperidin (C 12 F 23 N), have demonstrated high potential to treat tissues in hypoxic environments in cardiovascular tissue engineering, tumor treatment, and surgical applications. However, these treatments are not approved by the Food and Drug Administration (FDA) and are currently in clinical trial stages 67 . Table 3 summarizes the most common forms of fluorinated compounds and their applications as oxygen carriers.
The molecular structures of these surfactants, such as the fluorocarbon/hydrocarbon amphiphile FnHm (i.e., C 6 F 13 C 10 H 21 ) enables them to have a self-assembling nature and form highly stable emulsions and gels. This feature allows greater control over diffusion and oxygenrelease kinetics. In applications where there are limitations in oxygen diffusion or oxygen, there is typically an overall low cell viability, especially in organ-size constructs 67 . Fluorinated emulsions have been demonstrated to support these oxygen demands and enhance cell viability. Some fluorinated surfactants, such as perfuoromethyl-cyclohexylpiperidin, have shown high potential to treat tissues in hypoxic environments in cardiovascular tissue engineering, tumor treatment, and surgical applications 67 . However, these treatments are not approved by the Food and Drug Administration (FDA), which can delay clinical transition stages 67 . Table 3 summarizes the most common forms of fluorinated compounds and their applications as oxygen carriers. Fluorinated compounds, such as perfluorocarbons (PFCs), can dissolve large amounts of oxygen and act as oxygen carriers. In hydrogels, PFCs have been utilized in the form of aqueous emulsions or encapsulated directly in polymeric biomaterials. This incorporation releases the dissolved oxygen through diffusion. PFCs are denser materials that are advantageous in tissue applications that require an oxygen-rich environment 11,68,69 .

Applications of oxygen-releasing biomaterials in tissue engineering
Different types of oxygen-releasing chemistries achieve distinct oxygen-release characteristics. The cells in the body have varying oxygen tensions and saturation requirements for their survival, differentiation, functionality and growth. This section discusses how different oxygen-releasing biomaterials are used in applications for cardiovascular tissue engineering, islet transplantation, wound healing, skin regeneration, muscle regeneration, intervertebral disc regeneration, and preventing biofilm formation.

Cardiovascular tissue engineering
The challenges faced in the clinical translation of cardiovascular tissue engineered approaches include the lack of an ideal cell population, the inability to engineer cardiac tissue mimetic matrices, the inability to induce proper vascularization, and limitations in vascularization [70][71][72] . These factors are particularly important for the generation of physiologically-relevant 3D constructs 27 . Ensuring the oxygenation of engineered scaffolds is crucial for tissue viability and integration in vivo 30 . Control over oxygen-release kinetics can be achieved through the encapsulation of oxygengenerating materials in hydrogels and other polymeric materials at different concentrations. Due to the highly reactive nature of calcium peroxide, monitoring the oxygen-release rate can determine viability and induce cardiac stem cell differentiation 49 . Long-term cardiac cell cultures require sustained oxygen generation, which has been explored more recently with the addition of hydrophobic barriers.
Free radicals are considered a potential source of cytotoxicity in many 3D-encapsulation strategies. These radicals generated from the hydrolysis of solid peroxides could damage cardiac cells 16 . As an alternative approach, liquid peroxides, such as hydrogen peroxide, have been used in some cardiac tissue engineering applications to improve cell viability and provide sustained oxygen release with hydrophobic barriers. Figure 2 shows the use of liquid peroxides with a hydrophobic barrier to fabricate biomaterials that provides sustained oxygen release over time 49 . In a study by Li Z et al., hydrogen peroxide and poly (2-vinlypyrridione) (PVP) were encapsulated in PLGA microparticles, which were then incorporated into a thermosensitive hydrogel made from hydroxyethyl methacrylate oligo (hydroxybutyrate), N-isopropylacrylamide (NIPAAm), and acrylic acid (AAC). Cardiospherederived cells (CDCs) were encapsulated in this hybrid hydrogel. The results demonstrated the homogeneous distribution of cells throughout the 3D structure. The cell viability significantly increased in the presence of the oxygen-generating hydrogels placed in a hypoxic environment for up to two weeks (Fig. 2) 18,49 . Figure 3 shows another study in which primary human coronary artery smooth muscle cells (HCASMCs) were seeded onto scaffolds composed fluorinated zeolite (FZ) particles embedded in poly(carbonate-urethane) (PCU) (Fig. 3) 73 . These zeolite particles acted as oxygen vectors enabling enhanced oxygen delivery to the seeded cells. Modified and reprinted with permission from Reiss et al. 68 Copyright 1998, Elsevier The PCU-FZ scaffolds were coated with fibronectin for cell attachment and cultured for different time points: 4, 7, and 14 days 73 . Significantly higher cell proliferation and infiltration depths were observed in the PCU-FZ scaffolds than in the non-FZ PCU scaffolds 73 . Unlike blood, where oxygen is chemically bonded to hemoglobin, soluble oxygen can be released from the perfluorinated molecules whenever required. Figure 3 represents the viability and growth of the HCASMCs on PCU, PCU-FZ, and non-FZ PCU scaffolds. The results indicated that the PCU-FZ scaffolds supported higher cell proliferation than did the pure PCU scaffolds. The oxygen release increased as the fluorinated zeolite content increased. Based on the results of these studies, peroxides and fluorinated compounds incorporated into polymeric materials can enhance cell viability, growth, proliferation, and differentiation for cardiovascular tissue engineering applications.

Pancreatic tissue engineering
The transplantation of isolated islets of Langerhans is a potential treatment option for Type 1 Diabetes Mellitus (T1D) 74 . Numerous strategies have sought methods to induce angiogenesis after transplantation 75 . However, the introduction of angiogenic growth factors produced by the transplanted islets is insufficient to sustain the high oxygen demand immediately postimplantation. The existing approaches to maintaining the viability of transplanted islets use external oxygen supplies. Specifically, to suffice this demand, the patient is injected with oxygen delivery probes. The 3D tissue construct includes proangiogenic factors, such as vascular endothelial growth factor (VEGF). Similarly, prevascularization of the engineered constructs is also implemented to support tissue construct viability and oxygen demands 74 . Nevertheless, these current strategies have limitations in the efficiency  49 Copyright 2012, Elsevier of growth factor incorporation and do not guarantee high cell viability and functionality 76 . The development of oxygen-releasing scaffolds that improve tissue viability and functionality after transplantation has been explored and has shown promising results. In one study, oxygengenerating microparticles (MP) were engineered using a core-shell approach. The core comprised PVP/H 2 O 2 , which was formed inside a PLGA shell (Fig. 4) 77 . Figure 4 shows the structure of the microparticle surfaces, which were further coated with the enzyme catalase to ensure high cell viability. In the presence of catalase, hydrogen peroxide decomposed to oxygen and water after its release from the microparticle core. The microparticles were embedded in a nanofibrous scaffold made of collagen with a fibrin heparin/VEGF frame acting as the delivery system. High oxygen release between 0.83 ± 0.15 and 0.06 ± 0.05 mg/L was observed over a 14-day period 77 . The viability of isolated rat islets was monitored under both hypoxic and oxygen-generating conditions using MTS and LDH activity assays (Fig. 4). In the presence of the oxygen-generating microparticles, there was an increase in cell function, and cell lysis was not present 77 .
Solid oxygen-releasing compounds, such as sodium percarbonate (SPO) and calcium peroxide (CPO), have also been utilized to provide supplemental oxygen to pancreatic islets. In this approach, SPO particles were incorporated into polydimethylsiloxane (PDMS) in Petri dishes to generate 1 mm-thick films. Then, another 1 mm-thick layer of PDMS was poured and cured on top of the SPO film. Next, islet cells were seeded onto these oxygen-generating polymer layers to provide additional oxygen for 4 days 45 . Similarly, CPO particles were coencapsulated inside alginate microspheres. The CPO particles released oxygen via hydrolytic decomposition and enhanced islet viability. MTS assays indicated doubled cell viability in these oxygen-generating biomaterials 45 . Similarly, a composite oxygen-generating polymer was also synthesized to increase oxygen availability during the early islet engraftment period. Figure 5 shows the material consisting of a CPO encapsulating PDMS composite. Mouse insulinoma 6 (MIN6) cells were cultured overnight in a hypoxic environment with and without the PDMS-CPO discs 50 . The viability and metabolic activity of the MIN6 beta cells were higher in the presence of the oxygen-generating PDMS-CPO discs than in the absence of CPO. Figure 5 shows the oxygen-release kinetics observed in the in vitro culture of the MIN-6 cells. The positive effect on metabolic activity included sustained oxygen release over a 40-day period. These results were shown through an MTT assay conducted over a 25-day period. Overall, the results indicate that the long-term viability and proliferation of the beta cells after 3 weeks in  73 Copyright 2011, Elsevier cell culture were substantially higher on substrates with the oxygen-generating polymers than in control substrates where the oxygen-releasing compounds were absent (Fig. 5).

Wound healing
Wounds are common forms of injury in tissues with most acute wounds healing properly without medical interventions. However, some of the chronic wounds will have delayed healing or lead to unhealed open wounds or ulceration 78,79 . The blood supply is an essential component in wound healing that aids in the formation of connective tissue. In some cases, the local blood supply is limited, which inhibits the migration of inflammatory cells and proliferation of endothelial cells in wounds 80 . The formation of vasculature, however, is an extensive process in which injured tissue may be highly susceptible to necrosis 81 . Wound oxygenation treatment is a strategy to prevent necrosis and facilitate proper wound healing. Oxidation therapy has proven to be effective for hypoxic tissues 82 . Oxygen can improve the proliferation of fibroblasts by inducing these cells to produce more collagen, which enhances the formation of connective tissue 79,83 . Conventional therapies, such as hyperbaric oxygen therapy, have been used to treat a variety of open wound types, including diabetic ulcers. In combination with other current treatments, these methods in medical practice have shown therapeutic benefits 69,[84][85][86][87] . Oxygenreleasing materials have been employed in wound healing applications in the form of fluorinated methacrylamide chitosan and biodegradable hydrogels 79 . In this study, methacrylamide chitosan modified with perfluorocarbon chains (MACF) hydrogels were placed in phosphatebuffered saline (PBS) and saturated with up to 100% oxygen for 10 min Then, these hydrogels were placed on dissolved oxygen sensors in a closed container. Oxygen saturation was measured at three time points (1, 24, and 48 h). The results showed that the hydrogels could effectively deliver and release oxygen in vitro 79 . The hydrogel was then implanted onto a rat skin wound for 8 days. Histology and immunohistochemistry analyses demonstrated an enhanced wound healing process in the presence of the oxygen-releasing hydrogel 79 .
In another study, a cost-effective and portable device was developed to release oxygen to a targeted damaged tissue area without affecting other undamaged tissues 78 . The oxygen delivery device consisted of parchment paper and PDMS, on which microfluidic channels were fabricated. Hydrogen peroxide was introduced in the PDMS channels to release oxygen to the paper, which contained manganese dioxide embedded in catalytic regions. The microfluidic channels that included the oxygen-releasing component released oxygen at a controlled rate. The results of the in vitro study showed high cell viability in the presence of oxygen-releasing reagent. In some cases, hypoxic wounds can develop into ulcerations that do not heal efficiently. Diabetes is the most common reason for foot ulcerations, which can directly impact the mobility and overall quality of life of patients 88 . In another study, PVP and polyvinyl alcohol (PVA) hydrogels were prepared as oxygen-releasing materials 89 . PVP and PVA were crosslinked using gamma-ray irradiation, in which glucose oxidase and peroxidase enzymes were immobilized to generate oxygen. Materials that provide well-oxygenated environments support proper wound healing. Necrosis is a common issue associated with biomaterials in the treatment of chronic wounds. Polymer oxygen-generating (POG) films have been used to minimize necrosis by incorporating SPO into films of PLGA (Fig. 6). Figure 6 shows the results from a study that involved oxygenreleasing polymer films to reduce hypoxia in induced necrosis and support wound healing. In this study, rectangular flaps were created on the backs of nude mice, which were categorized into two groups 51 . One group received pure PLGA films, and the other group received   POG films. Flap necrosis was significantly less in the POG film-treated wounds than in the pure PLGA films (Fig. 6) 51 . Based on these findings, these technologies are advantageous and highly translatable in a series of wound healing applications.

Bone tissue engineering
The development of ideal barrier membranes with appropriate porosity and bioactivity is essential for the guided growth of bone in orthopedic and craniomaxillofacial applications [90][91][92] . A bone tissue engineering scaffold must be osteoconductive, porous, and biodegradable to support the attachment and proliferation of bone cells and encourage bone formation [93][94][95] . Furthermore, the blood supply must be compatible with artificial scaffolds to promote new tissue growth and healing 96 . Recently, a direct-write assembly (robocasting) technique was introduced to create porous scaffolds for bone regeneration. Compared to conventional approaches (e.g., freeze drying, gas foaming and electrospinning), this technique enables easier fabrication of complex structures 97,98 . Figure 7 shows the results from a study in which a biphasic calcium phosphate scaffold was fabricated with 60% hydroxyapatite (HA) and 40% beta-tricalcium phosphate (β-TCP) (Fig. 7) 99 . The scaffold was then coated with polycaprolactone (PCL) and CPO incorporated. Once the scaffold was submerged in culture medium, oxygen was generated and released. The 10-day oxygen-release kinetics analysis showed no indication of a burst of oxygen (Fig. 7) 99 . Osteoblasts were then seeded onto the scaffold and cultured for up to 7 days. Staining with 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) and dialkylcarbocyanine demonstrated that the oxygenreleasing scaffolds did not damage the cells (Fig. 7).
Although oxygen supply is important, the uneven oxygen distribution within 3D scaffolds can create issues in the mineralization process during bone formation 100 . To study this aspect, oxygen levels were monitored for a static 3D culture system with mouse osteoblasts for 7 days. The results showed that the central part of the scaffold had significantly lower levels of oxygen compared to the outer parts of the scaffold. The findings of the study suggested that implementing oxygen-releasing scaffolds supported homogeneous oxygen tension. Therefore, scaffolds enriched with oxygen-releasing materials are expected to provide an environment to facilitate osteogenic differentiation of MSCs and other precursors, leading to homogeneous vascularization in vivo. These strategies can also benefit targeting hypoxia responses that can arise in bone tissue engineering or musculoskeletal regenerative applications

Muscle regeneration
There are a plethora of diseases and conditions that can cause muscle tissue damage or degeneration. Trauma or tumor formation often causes muscular tissue damage. Ischemia and ischemia-reperfusion injury negatively affect the recovery of damaged muscles and tissue regeneration 101 . Hyperbaric oxygen therapy has previously been used to treat these conditions; however, this approach was not an effective therapy 102 . An oxygen-releasing biomaterial, sodium percarbonate (SPO), has been used for skeletal muscle repair 103 . In this study, SPO particles were added at concentrations of 0, 1, and 10 mg/mL in Dulbecco's Modified Eagle's Medium (DMEM). The oxygen generation, hydrogen peroxide content, and the consequential pH changes were monitored. The dissolved oxygen concentration was determined by obtaining readings using a fluorescence plate reader instrument. SPO was tested in vitro for its oxygen-releasing properties and biocompatibility. In an aqueous environment, SPO was shown to be an in situ oxygen-releasing material with high biocompatibility. This oxygen-releasing reagent was then injected into ischemic muscles in rats, which resulted in significant improvements in limb muscle regeneration 103 . In a similar application, SPO particles were incorporated into electrospun nanofibers. Then, rat muscle precursor cells were cultured in these scaffolds under hypoxic environments in vitro. The oxygenreleasing nanofiber mesh yielded improved viability and differentiation 34,103 .
Calcium peroxide has also been studied as a particulate oxygen-generating (POG) material with applications in enhanced muscle regeneration. In this approach, SPO and CPO were used in combination to engineer oxygenreleasing materials for tissue regeneration 104 . Skeletal muscle tissue was selected due to its high oxygen demand and dependency. The primary muscle tissues were isolated from rats and used in an in vitro hypoxic skeletal muscle injury model. The function of the engineered muscle tissue was analyzed before and after POG injection. The researchers injected the POG materials into the injured muscles of rats in the in vivo model. The histology experiments were performed for the recovering tissue after 48 h. The results revealed that the injection of POG created an environment that was conducive to muscle regeneration and functionality 104 . These oxygen-releasing biomaterials have been proven to effectively stimulate muscle recovery and restore muscle function.

Intervertebral disc regeneration
Neck and lower back pain and discomfort are commonly associated with intervertebral disc degeneration. The current treatments, such as corticosteroid injections, surgeries, and cell-based therapies, have been shown to be ineffective. Most treatments are aimed at solely alleviating the pain instead of treating the degenerated disc. Similarly, most new cell-based strategies are in early stages and may not be compatible with all patients 105 . Recently, Fig. 7 Oxygen concentration and corresponding cell viability of osteoblasts on the scaffolds with varying concentrations of calcium peroxide (CPO). a Oxygen concentration in the control group and in the 1, 3, and 5% (w/v) CPO-coated PCL scaffolds. b Cell viability of the osteoblasts on the CPOcoated scaffolds on days 1, 3, 5, and 7. Higher cell viability was observed for 3% (w/v) CPO-coated scaffolds. (c) Fabrication process of a 3D printed scaffold for bone regeneration using 3D printing. d Laser confocal microscopy images of osteoblast cells on (i) an uncoated BCP scaffold, (ii), (iii), and (iv) CPO-PCL BCP scaffolds coated with 1, 3, and 5% CPO, respectively. Modified and reprinted with permission from Touri M. et al. 99 Copyright 2018, Elsevier oxygen-releasing biomaterials have been shown to regenerate intervertebral discs. Inadequate nutrient and oxygen supply to disc cells, or nucleus pulposus (NP) cells, is one of the main causes of disc degeneration. Lack of oxygen prevents NP cells from synthesizing and preserving the extracellular cell matrix in the discs, leading to disc degeneration 105 . To address this problem, alginatebased scaffolds were prepared in combination with different concentrations (2.5%, 5%, and 10% w/v) of perfluorotributylamine (PFTBA) to reduce disc degeneration 106 . Using PFTBA as an oxygen regulator, in vitro and in vivo studies were conducted. Figure 8 shows in vitro cell culture studies under hypoxic conditions produced an increase in the synthesis of ECM components (aggrecan, collagen II) for NP cells. Increased GAG expression in these cells occurred in 2.5% (w/v) PFTBA-alginate scaffolds. The expression of these components was observed in normal NP cells, and the loss of their expression was observed in disc degeneration. Moreover, PFTBA at a 2.5% (w/v) concentration supported NP cell survival and proliferation 106 . An intervertebral disc degeneration model was used in mice for the in vivo experiments. The mice that were injected with 2.5% (w/v) PFTBA-alginate scaffolds and NP mouse cells showed alleviated disc degeneration. Specifically, the scaffolds supported the restoration of disc height and extracellular matrix components. These results were determined based on the disc high index (DHI) analyzed through micro-CT scanning and immunofluorescence studies for aggrecan and collagen II expression (Fig. 8) 106 .

Other tissue engineering applications
With a myriad of highly vascularized tissues, oxygenreleasing biomaterials are highly applicable in treating many forms of diseased or damaged tissues. For instance, perfluorocarbon-infused oxygen-generating polymers are strategies used in treating highly sensitive hypoxic tumors 107 . Oxygen-releasing biomaterials have also been used in neural stem cell engineering. Perfluorocarbons in conjugation with methacrylamide chitosan were found to promote the differentiation and growth of neural stem/ progenitor cells 108 . In plastic surgery, oxygen-generating polymer microspheres have been previously utilized to improve the efficiency of fat transplantation 109 . Autologous transplantation of fat is common in plastic surgeries, but a high rate of resorption has prevented this technique from practice until the introduction of oxygenreleasing materials. Adipose-derived stem cells (ASCs) have been used with oxygen-releasing microspheres for the transplantation of fat. The oxygen-releasing microspheres supplied oxygen to the ASCs where diffusion limitations interfered with adequate levels of oxygen. The in vivo results in rat models of this study demonstrated a significantly higher survival rate 2 weeks at postoperation.
The ASCs that survived were metabolically active in adipocyte angiogenesis, which improved the success of the fat transplantation 109 . Another tissue that is metabolically active with high oxygen demands is the liver. Hepatocytes of the liver are metabolically active cells but have limitations in oxygen diffusion in organ-sized constructs 110 . The incorporation of fluorocarbon-based oxygen-releasing systems has proven to effectively increase cell metabolic activity and proliferation 110 . For example, a collagen matrix included perflubron as an oxygen-release component 110 . The results of this study demonstrated an increase in the oxygen concentration in tissue cultures up to six times. These applications demonstrate that oxygenreleasing biomaterials have exceptional value for the repair and regeneration of tissues and can potentially bring tissue engineering approaches closer to clinical translation.

Preventing biofilm formation
Infections are one of the primary causes of biofilm formation on tissues surrounding implants in vivo. Consequently, biofilms produce an unfavorable environment for implant-tissue integration and require the implant to be removed and replaced 111 . Recent approaches to prevent biofilm formation use methods that involve modifying the surface of these implants with polymers, such as polyethylene glycol (PEG), which have anti-adhesive and antibacterial properties. Other approaches include coating the surfaces of implants with positively charged antimicrobial polymers and incorporating bleaching agents 112 . However, these modifications result in poor tissue integration in vivo. Furthermore, the long-term in vivo use of the implants leads to adsorption of plasma proteins on the surface of the implants. This process causes the inactivation of coated antimicrobial properties. Thus, these modifications are tailored for short-term applications 113 . Novel approaches must involve biomaterials that eliminate biofilm formation and support satisfactory implant-tissue integration for long-term application. Recent studies employ oxygen-releasing biomaterials, of which the byproducts (e.g., oxygen radicals) act as antimicrobial agents with low to no toxicity. An electrospinning method was used to create oxygenreleasing PCL nanofibers with different concentrations of calcium peroxide (1, 5, and 10% w/v) 113 .
In addition, ascorbic acid has been incorporated into these nanofibers to eliminate the toxic effects of CPO byproducts and support human osteoblast (hFOB 1. 19) viability. The antimicrobial properties of the oxygenreleasing nanofibers were evaluated by an MTS assay, in which oxygen-releasing nanofibers (10% (w/v) calcium peroxide-containing PCL nanofibers) were incubated with Escherichia coli (E. coli) or Staphylococcus epidermidis (S. epidermidis) strains. A Kirby-Bauer test was performed for E. coli growing on an agar plate to evaluate the microbial activity of the oxygen-releasing PCL nanofibers. The results demonstrated that a 10% (w/v) calcium peroxide addition in the PCL nanofibers eliminated E. coli and S. epidermidis growth by 95 and 90%, respectively. An inhibition zone formed around the antibiotic nanofibers that had been placed on the agar plates. The 2D seeding of human osteoblasts on calcium peroxide-PCL nanofibers resulted in short-term toxicity during the first day of incubation. This toxicity resulted from the initial rapid release of CPO from the nanofibers. However, the viability of the osteoblasts increased after four days of incubation. Greater viability was observed in the condition involving the integration of ascorbic acid within the calcium peroxide nanofibers. In orthopedic implants, the application of such nanofibers can prevent the bacterial formation and promote integration into the existing bone tissue. Fig. 8 An in vivo study for PFTBA-alginate-based scaffolds in a mouse disc degeneration model. a Represents the number of viable cells on days 0, 2, 4, 6, and 8 and metabolic activity recorded as absorbance under: normoxic, hypoxic, hypoxic +2.5% PTFA, +5% PFTA, +10% PTFA respectively. Control group (b), 2.5% (w/v) PFTBA group (c), 5% (w/v) PFTBA group (d), 10% (w/v) PFTBA group (e), alginate group (f), and puncture group (g) represent the images obtained from the micro-CT evaluation at 6 weeks post-operation. Modified and reprinted with permission from Sun Z. et al. 106 Copyright 2016, Elsevier

Conclusions and future directions
Tissue engineering approaches have been successful in formulating strategies to support the formation of vasculature in 3D constructs for the repair and regeneration of tissues. The synthesis of oxygen-releasing materials enhances these modern methodologies and addresses the high oxygen demands in tissue constructs. Oxygenreleasing biomaterials in the 3D encapsulation of cells offers an environment that meets the high oxygen demand posttransplantation in vivo. Many oxygenreleasing materials incorporate solid peroxides, liquid peroxides, and fluorinated compounds that decompose and release oxygen 114 .
Among many studies, calcium peroxide has been demonstrated to be the most effective inorganic solid compound for oxygen generation 59 . As discussed, for a controlled rate of reaction, calcium peroxide can be incorporated into hydrophobic polymers. The controlled and sustained presence of oxygen is imperative for the first few weeks to promote cell differentiation and tissue healing 52,115 . The use of liquid peroxides (e.g., hydrogen peroxide) as oxygen carriers has also shown considerable potential in tissue engineering applications. Specifically, the high solubility of hydrogen peroxide can improve the oxygen-carrying capacity of a polymer 11 . In addition, fluorinated compounds, such as perfluorocarbons, have shown promising performances as oxygen carriers and in supporting cell viability 116 . Overall, oxygen-releasing biomaterials have been widely explored in cardiovascular tissue engineering, islet transplantation, bone regeneration, muscle regeneration, and wound healing applications 51 . This vast array of fields also includes intervertebral disc regeneration 106 and preventing the formation of biofilms 113 .
The other prospects for oxygen-releasing polymers include drug delivery, preventing necrotic cell death, and tissue rehabilitation. Oxygen-releasing biomaterials can offer therapeutic and functional benefits during tissue regeneration and remodeling. In synthetically bioengineered implants, vascularization and tissue ingrowth are necessary for integration in vivo. The studies in this review have revealed that an oxygen-rich environment can encourage integration with the host tissue. This particular aspect is applicable for tissue integration around prostheses and orthotic implants. Moreover, the applications support sufficient nutrient and oxygen diffusion in vivo for viable and functional tissue constructs. In further developments, systematic in vivo studies and histological analyses are necessary to validate these specific strategies. In these biomaterials, controlled oxygen release is also applicable for drug delivery processes 117 . Based on previous applications, hydrophobic barriers can assist in facilitating the controlled and targeted delivery of pharmaceutical products. The versatility of oxygen-releasing biomaterials can give rise to next-generation viable and functional tissue constructs. As this area of research continues to grow, oxygen-releasing biomaterials can advance the translation of engineered constructs from the lab bench to patient care.