University of Groningen
Oxygen-releasing biomaterials
Steg, Hilde
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Steg, H. (2018). Oxygen-releasing biomaterials. Rijksuniversiteit Groningen.
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Oxygen-releasing biomaterials
Oxygen releasing biomaterials
ISBN 978 94 034 1033 3 (printed version) ISBN 978 94 034 1032 6 (electronic version)
Oxygen-releasing
biomaterials
Proefschrift
ter verkrijging van de graad van doctor aan de
Rijksuniversiteit Groningen
op gezag van de
rector magnificus prof. dr. E. Sterken
en volgens besluit van het College voor Promoties.
De openbare verdediging zal plaatsvinden op
dinsdag 6 november 2018 om 11.00 uur
door
Hilde Steg
geboren op 21 november 1984
te Lelystad
Promotores Prof. dr. D.W. Grijpma Prof. dr. S.K. Bulstra Copromotor Dr. R. Kuijer Beoordelingscommissie Prof. dr. R.R.M. Bos Prof. dr. R.A. Bank Prof. dr. D. Stamatialis
Contents
1
2
3
7
5
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4
IntroductionReview on increasing the viability of seeded cells in implanted tissue engineering scaffolds
Control of oxygen release from peroxides using polymers
Oxygen-releasing poly(trimethylene carbonate) microspheres for tissue engineering
Biocompatibility and proof of concept of novel oxygen delivering microspheres
Culturing of hMSC and SaOs-2 cells on oxygen releasing poly(trimethylene carbonate) and CaO2 composites under hypoxic conditions
General discussion
Summary / Samenvatting Dankwoord Curriculum vitae
1
Introduction
9
1
The loss or failure of bone is one of the most frequent, devastating, and costly problems in human health care. In 2005 traumatic bone fractures accounted for 8.5 million physician visits. Almost 1 million affected people required hospitalization1. One of the major problems in treating patients with
large bone defects is the healing of a very large bone defect which cannot be healed by the body itself; the non-union bone defect.
Physicians have treated patients with these problems for years, with varying results. Nowadays, most non-union bone defects are treated with tissue transplants. A bone tissue transplant can either be from autogenic, allogenic2 or
xenogeneic origin3. Each option has its advantages and disadvantages. Autogenic
tissue often results in donor site morbidity4, which is a large problem for the
patient. Moreover, the amount of bone that can be harvested, is limited. Allogenic and xenogeneic bone tissue do not have these difficulties, but its use can result in immune reactions towards the foreign body5 and carry the risk of
transmission of disease6,7.
In 1993 Langer and Vacanti introduced the concept of tissue engineering, a new field of research that aims at applying the principles of biology and engineering to the development of functional substitutes for damaged tissue8
(Figure 1). In tissue engineering, autogenic cells are isolated from the body, expanded in number in a cell culture laboratory and seeded in a suitable scaffold. Then the cell-scaffold construct can either be cultured further in the presence of growth- or morphogenetic factors that steer proliferation or differentiation of the cells. Finally, the cell-scaffold construct is transplanted at the site of the tissue defect to integrate with the natural tissue.
Introduction
Figure 19: A schematic overview of tissue
engineering:
1: Specific suitable cells are harvested from the patient’s body;
2: These cells are increased in number in vitro; 3: The cultured cells are seeded in porous scaffolds together with growth factors to stimulate their proliferation and/or differentiation;
4: The cell-seeded scaffolds are placed in culture to further increase cell numbers, or to produce tissue;
5: The tissue generated in the laboratory is then implanted at the site of damage to integrate with the natural tissue.
With tissue engineering, repair of large tissue defects in the human body would be possible without the drawbacks of allogeneic or xenogeneic donor tissue, since the regenerating tissue originates from the patient’s own cells. Although this technique seems very promising, 23 years after publication of the Langer and Vacanti’s paper, tissue engineering is still scarcely used in clinical settings10. An important problem in tissue engineering, is the fact that
immediately after implantation of a clinically relevant-sized construct the seeded cells die, likely due to the lack of oxygen and nutrients to the cells in the cell-scaffold construct9. Oxygen and nutrients are transported in blood, and the
necessary vasculature in a cell-scaffold construct is missing.
During cell culturing, nutrient and oxygen levels can be carefully monitored in the culture medium. However, a problem arises when the cell-scaffold construct is transplanted into the patient’s body. The cells in the construct will experience a hostile environment as the tissue is inflamed due to the wound healing reaction. The surgeon has disrupted the local vasculature capillaries11 and
11
1
attract both immune cells and stem cells to the wounded area and stimulate proliferation and vascularization15. Although this is a highly potent effect, the cells
that are used, are selected for their potential to regenerate, not for their potential in the trophic effect, a more easily accessible cell like blood platelets16,
might be a better source of trophic factors.
When cells are chosen for their capacity to either differentiate into a specific tissue or to proliferate and divide into cell types needed to restore the tissue of purpose, it will be essential to maintain the viability of the cells. Since the micro-vascularization is disrupted during the surgical procedure and the ingrowth of new vasculature is slow, the supply of oxygen, nutrients and the removal of waste products should be ensured in another manner. In the studies described in this thesis, we have focused on supplying oxygen to the cells.
Improved angiogenesis has shown to enhance the survival of transplanted cells to some extent, but only at the edges of the cell-scaffold constructs17. In in
vitro generated tissue, it has thus far not been possible to co-culture a vascular
system within the constructs. Skin, bladder tissue, cartilage and cornea5,18 can be
engineered without a vascular system, and only these tissues are currently applied clinically19. In vascularized tissues the limit for oxygen diffusion is 100–200µm 18,20,
and cultured cell-scaffold constructs have therefore been limited to these dimensions21. For a vascularized tissue such as bone, the diffusion of oxygen into
a cell-scaffold construct of clinically relevant size will be insufficient and the cells in the inner parts of the construct will die by necrosis due to hypoxia9.
To improve the viability of tissue engineered constructs upon implantation, different strategies have already been investigated. The porosity of scaffolds has been improved22 to enhance the diffusion of oxygen and nutrients into the
scaffold. Target cells have been co-cultured with endothelial cells to enhance the formation of new blood vessels10 upon implantation. Also a wide variety of
bioreactors, which include using hollow semi-permeable membranes as pseudo blood vessels23, has been developed. However, none of these strategies has led to
Introduction
Oxygen-releasing biomaterials
A scaffold prepared from a biomaterial that releases oxygen, could allow the cells in the construct to deal with the lack of oxygen until new vasculature has formed. For clinically successful cell-scaffold constructs, good control of the amount of oxygen released and the rate at which it is released is essential. When cells are provided with physiological concentrations of oxygen, sufficient to be normal metabolically active, the cells will not receive the trigger to express angiogenic growth factors24. However, when the amount of oxygen released is
too low or the time during which it is released is too short, the cells will die or go into a dormant state and not produce angiogenic growth factors25. In both
cases the healing process will not be completed, because the implant will not be vascularized. Oxygen-releasing scaffolding materials should supply the cells with just enough oxygen to preserve their metabolic activity while also produce angiogenic growth factors. Angiogenesis is a slow process: arteries develop by 100-200µm per day26,27 and several days to weeks (depending on the size of the
cell-scaffold construct) of oxygen-release are necessary to keep the implant viable.
Although research has emphasized the importance of physiologically relevant oxygen levels in situ28, too high oxygen levels should be avoided. Cell-based tissue
engineering is mainly based on the potential of mesenchymal- or hematopoietic stem cells to differentiate into cell lineages of choice. These differentiation processes are influenced by local oxygen tension and are more negatively affected in the presence of high oxygen concentrations29. A low oxygen concentration,
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water33–37. Peroxides react with water to produce oxygen, for example calcium
peroxide:
𝐶𝑎𝑂
2+ 2𝐻
2𝑂 ⇌ 𝐶𝑎(𝑂𝐻)
2+ 𝐻
2𝑂
2[Equation 1]
2𝐻
2𝑂
2⇌ 2𝐻
2𝑂 + 𝑂
2[Equation 2]
Other inorganic peroxides that release oxygen upon contact with water are MgO2, sodium percarbonate (2Na2CO3·3H2O2) and SrO2. The use of CaO2 is
preferred, as it has the lowest oxygen-formation rate38 and the simultaneously
formed Ca(OH)239 appears to induce bone growth.
By embedding the peroxide particles in a hydrophobic polymer matrix, an oxygen-release profile with a reduced initial release rate and a prolonged release duration can be achieved when compared to the oxygen-release characteristics from the particles alone. In earlier work, the suitability of using polymers as a controlling barrier between an active ingredient and the surroundings40–44 and the
release of oxygen from composite biomaterials45,40,41,46 was shown. Examples of
biodegradable hydrophobic polymers that have been used in the preparation of such composite materials are: poly (lactide-co-glycolide) (PLGA), poly (DL-lactide) (PDLLA), poly(-caprolactone) (PCL) and poly(dimethylsiloxane) (PDMS). It should be noted that of these polymers PDMS is non-degradable47 and PCL only
degrades in the body at a very slow rate (≥ 1 year)48.
We further hypothesized that the inflow of water into the polymer would determine the rate of the oxygen-release from the composite. Therefore, the focus was to find a biodegradable polymer with a limited ability to absorb water.
Poly(trimethylenecarbonate) (PTMC) is a biodegradable polymer that has different interesting properties. The degradation behaviour of this biodegradable biomaterial is in a surface eroding manner mediated in vivo by the activity of macrophages. In vitro this surface erosion can be modelled using cholesterol esterase or lipase. Degradation of the material results also in non-acidic products, thereby the bulk release of acidic monomers, seen in the degradation of lactic acids49,50is overcome.
Introduction
In our approach, we have developed and evaluated the properties and performance of oxygen-releasing biomaterials by preparing composite materials based on biodegradable hydrophobic polymers and calcium peroxide.
Conclusions
A formidable challenge in tissue engineering is to prevent the seeded cells or newly generated tissue from dying shortly after implantation. Despite years of extensive scientific research, this problem has not been solved. Oxygen-releasing biomaterials have been designed, manufactured and tested and show some potential in preventing or postponing cell death. From these data, we can conclude that oxygen-releasing materials are worth investigating further.
The goal of this project was the production of a biocompatible material that can release low amounts of oxygen for up to 3 weeks. One of the sub-aims of the studies described in this thesis was also to investigate which factors are important for a prolonged release of oxygen, and which hydrophobic polymer is most suited as oxygen-releasing biomaterial. Finally, the way in which the material is applied, was subject of investigation.
Aims and structure of this thesis
The work described in this thesis aims at developing a functional biomaterial that releases oxygen over a prolonged period of time, thereby improving the viability of cells in vitro and also in vivo. Furthermore, to create a functional in vitro model to study the different effects of ischaemia and the effectivity of an oxygen-releasing biomaterial.
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In chapter 3 we aimed to produce microspheres from a slow oxygen-releasing composite using PTMC as a carrier material combined with CaO2. It
was demonstrated that this oxygen-releasing composite showed slow releasing properties combined with good cell compatibility. The microspheres of this material created an oxygen-releasing product which is easy to dose and can be added to scaffolds of other materials, such as ceramics.
In the study described in chapter 4 oxygen-releasing microspheres produced from PTMC and CaO2 composites were tested in vivo for their
functionality. The oxygen-release from the microspheres improved the viability of the otherwise ischaemic tissue.
The PTMC/CaO2 composite was also studied further in chapter 5 to
create a functional in vitro hypoxic model for a better understanding of oxygen-delivering biomaterials. Although the composites showed already effectiveness in the in vivo models, the model based on the absence or presence of oxygen in a cell culture set-up showed not to be enough to study the effect of the oxygen-delivering biomaterial. This indicates that the effectiveness of the biomaterial should be studied in a more complicated model and a small study was added towards different factors in an ischaemic system.
Introduction
References
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43. Bezemer, J. M., Grijpma, D. W., Dijkstra, P. J., van Blitterswijk, C. A. & Feijen, J. Control of protein delivery from amphiphilic poly(ether ester) multiblock copolymers by varying their water content using emulsification techniques. J. Control. Release 66, 307–20 (2000). 44. Uchida, T., Yagi, A., Oda, Y. & Goto, S. Microencapsulation of ovalbumin in
poly(lactide-co-glycolide) by an oil-in-oil (o/o) solvent evaporation method. J. Microencapsul. 13, 509–18 (1996).
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46. Ng, S.-M., Choi, J.-Y., Han, H.-S., Huh, J.-S. & Lim, J. O. Novel microencapsulation of potential drugs with low molecular weight and high hydrophilicity: hydrogen peroxide as a candidate compound. Int. J. Pharm. 384, 120–7 (2010).
47. van Kooten, T. G., Whitesides, J. F. & von Recum, A. Influence of silicone (PDMS) surface texture on human skin fibroblast proliferation as determined by cell cycle analysis. J. Biomed. Mater. Res. 43, 1–14 (1998).
48. Bat, E., van Kooten, T. G., Feijen, J. & Grijpma, D. W. Macrophage-mediated erosion of gamma irradiated poly(trimethylene carbonate) films. Biomaterials 30, 3652–61 (2009). 49. Grizzi, I., Garreau, H., Li, S. & Vert, M. Hydrolytic degradation of devices based on
poly(DL-lactic acid) size-dependence. Biomaterials 16, 305–11 (1995).
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2
Review on increasing the viability
of seeded cells in implanted tissue
engineering scaffolds
Hilde Steg, Arina T Buizer, Willem Woudstra, Albert G Veldhuizen, Sjoerd K Bulstra, Dirk W Grijpma, Roel Kuijer
Manuscript in preparation
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
Abstract
Most tissues have limited regenerative capacity and repair through the formation of a fibrous scar, which impairs their function. To restore tissue defects with functional tissue, therapies involving the application of autologous, tissue-specific cells have long been considered to have promising prospects. However, thus far only a limited number of such therapies have reached clinical application. The major reason for the failure of the investigated applications is the difficulty to include a vascular system in the constructed tissue. Applications that have reached the clinic most often concern tissues that are non-vascular (cartilage) or poorly vascularized and thin (bladder). Attempts to promote vascularization, either before, in a 3D culture, or after implantation, have failed. The lack of oxygen and nutrients, as well as the hostile inflammatory environment immediately after implantation, in most applications result in necrosis of the newly synthesized tissue. One possible way to limit this necrotic process would be to apply oxygen to both the wound bed in which the newly synthesized tissue graft is placed, and to the graft itself. For this, a biomaterial which can deliver oxygen and serves as scaffold for the tissue producing cells, would be an option. Although, these materials have been developed, an optimal oxygen-releasing biomaterial has not been found yet. In this review, the different oxygen-releasing materials that have been investigated are reviewed, as well as the factors that influence the release of oxygen.
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Current treatment of large tissue defects
To a certain extent, the human body is capable of regeneration. When the damage is too large for the body to restore, donor tissue can be used for repair. This tissue can originate from the patient’s own body (autologous), from a human donor (allogeneic) or an animal (xeno-transplantation). Although all these tissues generally result in acceptable repair, they have their specific disadvantages. Autologous tissue is the most compatible with the patient’s body, but the amount of tissue that can be harvested is limited and donor site morbidity is a recurrent problem1. Allogeneic tissue is more readily available, but limited availability of
tissue and problems with compatibility resulting in tissue rejection, are common2,3. Moreover, the transmission of disease and the existence of a foreign
body reaction can be a danger. For xenogeneic transplants these risks are even more existent. Engineering of new tissue from autologous cells would circumvent the disadvantages of transplantation of donor tissue.
Tissue engineering
Tissue engineering is the field of research in which ways are explored to engineer new tissue using biomaterials, growth factors and (stem) cells. Both combinations of biomaterials, growth factors, biomaterials with cells and combinations of materials with growth factors and cells are being evaluated. Here, the focus is on applications in which autologous (stem) cells are being used. Autologous stem cells are very valuable tools in healing the human body, since they are capable of differentiation into different lineages. Therefore, they can become or form cells that are able to repair different tissues. Mesenchymal stem cells (MSC) can be harvested from bone marrow, but also adipose tissue is used as a cell source4. MSC are known to be able to differentiate into bone, cartilage,
fat and muscle5,6, under the influence of outside stimuli like growth factors and
surface adhesion7–9. To create a new functional tissue, human mesenchymal stem
cells (hMSC) should be differentiated into the right lineage, either by the supply of growth factors or by induction from the surroundings.
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
To create a three-dimensional (3D) tissue, the cells need some form of support. In the body the cells are supported by extracellular matrix. For cell-based therapies, a biomaterial scaffold is mostly used to provide stability. Biomaterials of different origin have been used, and include bioceramics10,
biopolymers11, and synthetic polymers12. In most applications, it is desirable for a
biomaterial scaffold to degrade over time in such a way that there are no remnants left in the body. Its function will be taken over by the newly formed extracellular matrix13,14. Apart from providing support, biomaterials have also
been known to be able to function as an initiator of differentiation of the stem cells15. Both tissue integrating and -differentiating properties of biomaterials are
known and used13,16–18. Moreover, in the manufacturing process of the
biomaterial, the shape of the biomaterial can effectively be altered. McBeath et al.8 showed that the topography of the attachment spot for the cells can be a
useful trigger for the cells to differentiate into the cell type of choice, and the surface chemistry of the biomaterial is known to influence cell behaviour by signalling through the integrin binding19. Furthermore, growth factors can be used
to direct the cells into the right differentiation lineage. The biomaterials can be used to deliver growth factors, and different biomaterials have shown to be suitable as a slow release system to prolong the delivery as well as to control the local concentration of growth factors. This minimizes the number of injections required and avoids peak concentrations of growth factors in the body20,21.
Challenges in cell-based therapies
Most clinically relevant sized three dimensional constructs containing cultured cells or tissues prove to be not viable after implantation22,23. Oxygen
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thickness of the implant, and therefore to the oxygen concentration at that place27. Radisic et al.28 developed a biomimetic culture system in which cardiac
fibroblasts and myocytes were co-cultured on scaffolds with an array of parallel channels that mimics the role of a capillary network. By increasing the amount of oxygen carrier, the oxygen concentration could be varied. It was found that increase in oxygen content enhanced cell density and DNA content, the amounts of cardiac proteins, and their contractile properties28, supporting the hypothesis
that oxygen is of vital importance to an implanted tissue construct29. The fact that
oxygen-dependent cell- or tissue constructs become necrotic after implantation and that low oxygen dependent tissues have been reasonably successful, confirm that for larger 3D structures containing cells or tissue the availability of oxygen is important for a successful outcome in the clinic.
In the body, oxygen is normally delivered through erythrocytes which take up oxygen in the lungs and distribute the oxygen during their journey through the vascular network in exchange for carbon dioxide. A proper vascularization is essential for the vitality of our tissues. It is expected that vascularisation is also key for the larger cultured 3D structures30. A suitable biomaterial scaffold seeded
with autologous cells able to produce the required tissue, will not perform adequately when it is not vascularized. This immediately shows the challenge we are facing. Well-vascularized tissues like bone or muscle are built up from cells that cannot survive for a sufficiently long time without oxygen and nutrients, making cell-based therapies nearly impossible.
Tissues cultured in the absence of fluid flow are solely dependent on diffusion, which leads to the formation of a 100-200µm outer layer of tissue that receives oxygen by diffusion31. This thickness corresponds to the oxygen diffusion
limit in tissue, which is characteristically 100-200µm26. During cell culture,
oxygen- and nutrient levels can be carefully monitored in the culture medium. However, after transplantation of the cell-scaffold into a patient’s body, the cells in the construct will experience a hostile environment in which a wound healing reaction takes place that involves an inflammatory reaction32. The surgeon has
disrupted the local vasculature1 and the tissue contains high amounts of fibrin due
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
an anoxic environment. Additionally, the limited refreshment of tissue fluids results in acidic conditions, in the end resulting in cells dying by necrosis32,33.
It has been found that these dying cells can contribute to the healing process. Wu et al.34 showed that dying cells in a co-culture increase the growth
of the second cell type. Apparently dying cells release many growth- and chemotactic factors to the environment, resulting in a so-called trophic effect33.
Growth- and chemotactic factors produced by hMSCs, are beneficial for the healing process; these factors attract both immune cells and stem cells to the wounded area and stimulate proliferation and vascularization35. However, the
carefully harvested, isolated, and cultured cells were selected for their ability to build new, specialized tissue, and not for their ability to produce growth factors while dying. It is possible that the chosen cells are not the most suited cells to create a trophic effect. More easily accessible cells or cell components like blood platelets36, might be a better source of trophic factors.
Several research groups have investigated the simultaneous production of a vascular system during the production of the tissue-scaffold construct25,26. This
was found to be more difficult than expected. To date, a completely vascularized tissue has not been engineered26, and although different artificial vascularization
strategies, based on permeable membranes37,38, showed promising results when
implanted subcutaneously, until now these systems failed when evaluated in orthotopic implantation sites and are not ready to be used in clinical situations. Therefore, cell-based therapies are now facing a most challenging problem: how to maintain cells alive upon implantation when there is no vascularization in situ.
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tissue of purpose, it will be essential to maintain the viability if the cells. Since the micro-vascularization is disrupted during the surgical procedure and the ingrowth of new vasculature is slow, the supply of oxygen, nutrients and the removal of waste products should be ensured in another manner. Here we focus on supplying oxygen to the cells by means of an oxygen-delivering biomaterial.
Oxygen-delivering biomaterials
Based on the above-mentioned assumptions, a device able to release low but significant concentrations of oxygen for a period of 3-4 weeks throughout the newly synthesized tissue, would solve the difficulties in cell culturing mentioned above. If such a device would be the scaffold on which the cells were seeded before implantation, the seeded cells would no longer suffer from a lack of oxygen (Figure 1). The period of oxygen-release should suffice for the body to synthesize blood vessels throughout the entire implant. This will be necessary, since the delivery of oxygen as the sole nutrient is not enough for the cells or tissue to stay viable. Thereby, vascularization is also necessary for the removal of metabolites created by the viable cells.
Figure 1: A homogeneously seeded scaffold larger than 200µm3 will over time become hypoxic in
the centre, resulting in cell necrosis. A homogenously seeded oxygen-releasing scaffold will provide the oxygen necessary for the cells in the centre of the scaffold, thereby supporting cells during the period required for vascularization40.
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
Fortunately, by providing the cells with sufficient oxygen the amount of waste products will also decrease. In 1861 Louis Pasteur found that cells cultured in hypoxic or even anoxic environments start ‘fermenting’ glucose41. This
‘fermenting’ process, later called anaerobic glycolysis, is far less efficient than oxidative phosphorylation at normal oxygen concentration. Although anaerobic glycolysis is beneficial for short periods of hypoxia, after implantation of
engineered tissue it is unwanted because of resulting high concentrations of waste products like lactate in the tissue. Oxidative phosphorylation produces an 18-fold production of adenosine triphosphate (ATP) per glucose molecule compared to anaerobic glycolysis42 (Figure 2) and much lower concentrations of
waste products.
Figure 2: Oxygen is an important nutrient to produce energy. Although a cell is capable of anaerobic glycolysis when oxygen is not available, the amount of ATP produced is much lower42.
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optimal for cell viability. It was further hypothesized that the use of an oxygen-generating substance could decrease apoptosis and necrosis in newly formed cell-based tissues, but also provide a new therapy for other ischemic tissues like cardiac muscle after a myocardial infarct and chronic wounds like foot ulcers45,46.
In chronic wounds the pO2 is critical for the whole cycle from prevention,
towards treatment. Oxygen-delivery in situ can make the difference between amputation or a healthy limb40,47. Furthermore, Wang et al. showed the potential
for oxygen-releasing materials to inhibit bacterial infections48.
Key to an effective releasing biomaterial is the rate of oxygen-release. The release should be sustained and tuneable to be useful in different circumstances. Since an oxygen-releasing biomaterial will not release oxygen indefinitely, the tissue should be provided with enough oxygen to keep the cells viable, while on the other hand the oxygen concentration should be low enough to stimulate blood vessel formation. Some level of hypoxia is required to stabilize the hypoxia inducible factor-1 (HIF-1) complex in cells which leads to the formation angiogenesis-promoting growth factors such as vascular endothelial growth factor (VEGF) and fibroblast growth factor-2 (FGF-2)40.
Figure 3: Schematic overview of the effects of hypoxia40. Different hormones are upregulated by
hypoxia. The extent of hypoxia is very important: when hypoxic conditions become anoxic, the right part of the diagram becomes more dominant, resulting in dying cells.
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
Clark et al.39 studied the growth of endothelial sprouting in frog (Hyla
pickeringii) larvae and showed that the speed of ingrowing vessels is limited to tenths of micrometres a day. In bone lengthening studies in patients, the rate of blood vessel growth appeared to be similar. Ideally, the biomaterial would allow for the preparation of a tissue engineering scaffold that is larger than approximately 1cm3 and releases oxygen in a controlled manner for up to 21
days. The release of oxygen must be from several days to 3 weeks and allow for vascularization of a functional adequately-sized tissue. This should lead to a normal oxygen concentration in the tissue that ranges from 0.6–2.6kPa (4-20mmHg)41.
Oxygen-generating compounds
An essential substance in an releasing biomaterial is the oxygen-generating compound. Different oxygen-releasing materials have been explored. The use of substances originating from the body like encapsulated bovine haemoglobin in amylose to form a nano-sized oxygen carrier49,50 and chemical
blood replacers like perfluorocarbon (PFC) have been investigated. These compounds both have to be charged with oxygen before their application. Chemicals such as peroxides release oxygen upon contact with water and appear to be more suitable for sustained oxygen-release51–55.
A few oxygen-delivering biomaterials showing promising results have already been manufactured. Mostly, a solid peroxide like sodium percarbonate (2Na2CO3·3H2O2), calcium peroxide (CaO2) or magnesium peroxide (MgO2)
have been used as oxygen-releasing compounds. The oxygen-release rate from these peroxides depends on a large number of variables, including pH,
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The reaction mechanisms of the generation of oxygen from the solid peroxides described above include an intermediate step in which hydrogen peroxide is formed, hydrogen peroxide then forms oxygen:
CaO2 (s) + 2H2O Ca(OH)2 (s) + H2O2
MgO2 (s) + 2H2O Mg(OH)2 (s) + H2O2
(Na2CO3)2 • 3H2O2 4Na+ +2CO3- + 3H2O2
2H2O2 2H2O + O2
H2O2, CaO2, MgO2 and (Na2CO3)2•3H2O2 have already been used to
prepare oxygen-releasing biomaterials53,57–60,61. Although hydrogen peroxide
seems an optimal oxygen-generating compound since it leaves no waste product after the reaction, it is also highly reactive. It forms oxygen rapidly, which can lead to high concentrations of oxygen. Besides, proper encapsulation of hydrogen peroxide is essential to prevent its release and the sudden release of high concentrations of oxygen as both are detrimental to the cells.
Although a composite of sodium percarbonate and PLGA has been studied
in vitro58 and in vivo59, sodium percarbonate was not deemed suited as
oxygen-delivering component. Its high reactivity results in a high oxygen-release rates over a relatively short time period. Of the other metal-peroxides, CaO2 appears
most suitable based on its low water solubility and high purity62. Encapsulation of
CaO2 within a polymer matrix with limited swelling behaviour in water should
result in the sustained release of low amounts of oxygen. Its side reaction product, Ca(OH)2, is a substance that occurs naturally in bone and appears to
induce bone growth63.
Polymer matrix component of the
oxygen-delivering biomaterial
Another essential component for the slow release of oxygen is a polymer matrix to encapsulate the oxygen-generating compound. Encapsulation of the peroxide particles in a hydrophobic polymer matrix can be used to tune the oxygen-release profile to reduce the initial release rate and prolong the release duration. By encapsulation of peroxides in a hydrophobic, non-swelling polymer
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
matrix, the peroxides are shielded from body fluids. When the polymer matrix degrades or dissolves, the solid peroxide particles become exposed to water, the oxygen generating reaction starts and oxygen is released.
The intermediate step in the reaction of CaO2 with water leads to the
formation of Ca(OH)2 and H2O2. Since calcium hydroxide is a base with a low
solubility in water, the effect of pH on the reaction towards this intermediate will be evident. For this reason, the degradation mechanism, acidity and degradation speed will be critical for the formation of a sustained oxygen-releasing composite biomaterial. On the other hand, the influence of Ca(OH)2 on the pH in the body
will be limited due to its low solubility.
In earlier work, the suitability of using polymers as a controlling barrier between an active ingredient and its surroundings58,60,64–66, as well as the release
of oxygen from composite biomaterials48,58,60,67, were shown. Examples of
biodegradable polymers that have been used in the preparation of such composite materials are: poly(lactide-co-glycolide) (PLGA), in combination with sodium percarbonate58, hydrogen peroxide67 or calcium peroxide60,
poly(-caprolactone) (PCL) in combination with CaO248 and poly(dimethylsiloxane)
(PDMS) in combination with CaO257. It should be noted that of these polymers
PDMS is non-degradable68 and PCL only degrades in the body at a very slow rate
(≥ 1 year)69. The degradable biopolymers PCL and PLGA are non-soluble in
water, their degradation is based on the hydrolytic attack of the ester bond which leads to the formation of acidic compounds11. Furthermore, the
degradation of PLGA proceeds as a bulk degradation process70 resulting in a
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inorganic peroxide in a hydrophobic polymer matrix that has a very limited swelling capacity in water.
Our studies involved a biodegradable polymer with a limited ability to absorb water, that is biodegradable via a surface erosion process involving cellular enzymes like which cholesterol-esterase, does not lead to the formation of acidic degradation products69 (pH is known to affect the reaction of metal peroxides
with water56,71).
n Figure 4: structure of PTMC
The polymer of choice for this application is poly(trimethylene carbonate) (PTMC), it is a hydrophobic polymer with very interesting properties: it has an equilibrium water uptake of 1%, it is amorphous polymer, it is rubber-like at room- and body temperature and is biodegradable72 by surface erosion. In vitro,
cholesterol esterase can be used to induce surface erosion of the polymer. These properties make PTMC an interesting polymer for use as a matrix in oxygen-delivering composite biomaterials.
Conclusions
Cell-based therapies have been shown not to be less successful than expected. Implanted cells tend to die upon implantation because of the absence of an adequate vascularisation. Attempts to overcome the lack of vascularisation have thus far failed.
Oxygen-releasing biomaterials have shown to prolong cellular viability, but more research will be necessary to ensure their success. Most oxygen-releasing biomaterials have too short oxygen-release profiles or are non-biodegradable. To create a material that makes cell-based therapies possible, a slow oxygen-releasing biodegradable implant needs to be manufactured. Such materials and devices may not only influence cell-based therapies but may also help in treating and healing ulcer wounds and other problems related to ischaemia.
Review on increasing the viability of seeded cells in implanted tissue engineering scaffolds
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Control of oxygen-release from
peroxides using polymers
Hilde Steg, Arina T Buizer, Willem Woudstra, Albert G Veldhuizen, Sjoerd K Bulstra, Dirk W Grijpma, Roel Kuijer
Published in Journal of Materials Science: Materials in Medicine 2015; 26 (7) 207
Control of oxygen-release from peroxides using polymers
Abstract
An important limitation in cell therapy for the regeneration of tissue is the initial lack of oxygen. After implantation of large 3D cell-seeded structures, cells die rather than contribute to tissue regenerating. Oxygen-releasing materials were tested to improve cell survival and growth after implantation. Calcium peroxide (CaO2) in a polymer matrix was used as source of oxygen. Two
polymers were tested in order to slow down and extend the period of oxygen-release, poly(D,L-lactic acid) and poly(lactic-co-glycolic acid). Compared to CaO2
particles, both releasing systems showed an initially higher and shorter oxygen-release. Human mesenchymal stromal cells cultured on casted films of these oxygen-releasing composites required catalase to proliferate, indicating the production of cytotoxic hydrogen peroxide as intermediate. Poly(D,L-lactic acid) and poly(lactic-co-glycolic acid) are less suited for slowly oxygen-releasing materials. Catalase was able to reduce the cytotoxic effect of H2O2.
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Introduction
Cell therapy using autologous cells for replacing malfunctioning tissue is hampered by the lack of vasculature at the implantation site, resulting in cell death of the implanted cells1. Important factors causing cell death are considered
to be the limited amounts of oxygen and nutrients and the disability of cells to get rid of waste products. A potential solution for the lack of oxygen may be the use of composite oxygen-delivering scaffold materials, with peroxide salts as source of oxygen2. These materials should deliver oxygen to cells or tissue for a
prolonged period of time such that implanted cells survive and contribute to tissue repair, but without interfering with angiogenesis, which is induced in hypoxic conditions3–7. Such composites should provide cell survival throughout
the scaffold, thereby improving tissue restoration or repair. Oxygen-release should last until the angiogenic process is complete and a new functional vascular network is produced.
Peroxides provide oxygen by reaction with water (Equation 1). The reaction intermediate, hydrogen peroxide (Equation 2) is considered to be a cytotoxic agent. Mammalian cells have several defence mechanisms to help convert H2O2 into oxygen and water and can deal with low concentrations of
hydrogen peroxide7.
𝐶𝑎𝑂2+ 2𝐻2𝑂 ⇌ 𝐶𝑎(𝑂𝐻)2+ 𝐻2𝑂2
[Equation 1]
2𝐻2𝑂2 ⇌ 2𝐻2𝑂 + 𝑂2(Catalase)
[Equation 2]
A careful balance between oxygen delivery, angiogenesis and cytotoxicity is required. In this study oxygen-release from composite materials, consisting of CaO2 powder embedded in biodegradable poly-(lactic acid)(PLA) and
poly(lactic-co-glycolic acid)(PLGA) (Figure 1), were evaluated for their oxygen-delivering capacity and their cytotoxicity to human bone marrow mesenchymal stromal cells (hMSC). The polymer matrix was intended to act as a barrier for both inflow of water and outflow of active agent, slowing down the reaction and reducing cytotoxicity. Thus, we hypothesize that PLA and PLGA polymer barriers prolong oxygen-release and reduce cytotoxicity5,8,9. Since PLGA is faster
degrading than PLA, the PLGA-based composite is expected to show faster oxygen-release.
Control of oxygen-release from peroxides using polymers
Materials and Methods
Solutions of Poly(DL-lactic-co-glycolide) (MW 153,000 g/mol)(PLGA) and Poly(DL-Lactide) (PDL20 50/50, MW 400,000 g/mol)(PLA) (PURAC, Gorinchem, the Netherlands) in chloroform were prepared at 10%(w/v). Five percent (w/w) CaO2 (75% purity, 200 mesh, <75µm; Sigma-Aldrich, Zwijndrecht, Netherlands)
powder was dispersed and stirred vigorously. For oxygen measurements, 1ml of the suspension was pipetted in a 50mL vial and dried at room temperature for 48hours and in vacuum for another 24hours. Oxygen-delivery was measured at 37°C in 35mL deoxygenated Simulated Body Fluid(SBF)10 with a WTW Oxi 3310
oxygen meter (Weilheim, Germany). The oxygen measurement set-up was an open system, oxygen-free through a continuous in- and outflow of N2 gas.
Oxygen-release from CaO2 alone was assessed from 5mg CaO2 packed in a
porous filter paper placed at the bottom of the flask, which was then filled with 35mL SBF.
For cell culture, 15mm cover glasses were coated with 250µL suspension and dried as described above. Control materials were produced from 10% polymer solutions, not containing CaO2. hMSC, isolated from bone marrow of
patients receiving a total hip replacement (Buizer et al)11, were cultured in
alpha-MEM (Life technologies Europe, Bleiswijk, Netherlands) supplemented with 10% heat-inactivated Fetal Bovine Serum (FBS)(Life Technologies), 0.2mM Ascorbic-acid-2-phosphate (Sigma-Aldrich) and 1% antibiotic-antimycotic solution (Life Technologies) in an humidified atmosphere supplemented with 5% CO2. To
assess cytotoxicity, passage 3 hMSC 10,000 /well were seeded in a 24-wells plate containing PLA, PLA/CaO2, PLGA, PLGA/CaO2-coated coverglasses and cultured
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Statistical analyses were performed using a Univariate Anova and post-hoctests in SPSS 20.0.0.2. Time, material and the presence of catalase were the variables assessed.
Results and discussion
Oxygen-release of composite materials and calcium peroxide particles was determined in an open system, flushed with nitrogen at 37°C. Embedding of CaO2 in PLA and PLGA resulted in a faster release of oxygen compared to the
release from CaO2 crystals alone (Figure 2).
Figure 2: Oxygen-release in time from CaO2 and CaO2 crystals embedded in polylactide polymer.
Oxygen-release measurements were performed in an ‘open’, anoxic system at 37°C.
Embedding of CaO2 crystals in polydimethylenesiloxane (PDMS) to reduce
the amount of water influx and H2O2 outflow, resulted in a very effective
slow-release system for oxygen5. However, PDMS is a non-degradable material and has
been shown to have little cell adhesive properties5,12. Hydrolysis of the
lactide-based polymers may have resulted in lowering of the pH, thereby inducing a higher solubility of the intermediate Ca(OH)2, shifting the reaction towards H2O2
production13. 0 1 2 3 4 5 6 0 1 2 3 4 Ox yg en ( m g/ L) Time (days) PLGA CaO₂ P(DL)LA CaO₂ CaO₂
Control of oxygen-release from peroxides using polymers
PLGA-based composite materials for oxygen-release were previously used by Harrison4 and Oh3, either combined with Na2CO3·1.5H2O2 or with CaO2.
The observed oxygen-release kinetics of these studies are difficult to compare to our data, since these investigators used a closed system to assess oxygen-release and our setup was an open system more resembling an in vivo situation3.
Furthermore, the production method of the oxygen-delivering composite differs, which may have decreased the influx of water in their material as compared to ours. Our setup, an open system, did not allow for assessment of the total amount of oxygen delivered by the material.
A B
Figure 3: Histogram representing absorbance values of the XTT viability assay. Human mesenchymal stromal cells were seeded onto oxygen-releasing and non-oxygen-releasing materials and cultured in a hypoxic (0.1 %O2) environment. To some cultures catalase was added to the culture medium
to reduce the H2O2 concentration, thereby decreasing cytotoxicity. Cells were cultured for 1, 4, or
7 days. Statistical significant differences are shown using *. Furthermore, differences between with and without catalase are statistical significant for PLA/CaO2 and for PLGA/CaO2 at t=4 and t=7.
0 1 2 3 PLA PLA CaO₂ PLGA PLGA CaO₂ TCPS absorba nce ( 46 0 nm ) PLA PLA CaO₂ PLGA PLGA CaO₂ TCPS with catalase t=1 t=4 t=7 * [ * [ * [ * [ * [ * [ * [ * [ * [
