Supercritical fluid technologies and tissue engineering scaffolds

Current Opinion in Solid State and Materials Science 8 (2004) 313–321

Supercritical fluid technologies and tissue engineering scaffolds Robin A. Quirk a, Richard M. France a, Kevin M. Shakesheff b

b,* ,

Steven M. Howdle

c

a RegenTec Ltd, BioCity Nottingham, Pennyfoot Street, Nottingham NG1 1GF, UK School of Pharmaceutical Sciences, The University of Nottingham, University Park, Nottingham NG7 2RD, UK c School of Chemistry, The University of Nottingham, University Park, Nottingham NG7 2RD, UK

Received 23 October 2003; accepted 18 December 2003

Abstract Supercritical fluid (SCF) processing methods possess advantages over standard processing methods for the production of scaffolds for use in tissue engineering. Advantages include the absence of organic solvents, the ability to incorporate delicate biologicals without loss of activity, and control over the morphology of an internal porous architecture. This review describes SCF processing methods of relevance to tissue engineering and controlled release strategies, with focus on the incorporation of bioactives such as protein growth factors. Ó 2004 Elsevier Ltd. All rights reserved. Keywords: Supercritical fluid; Carbon dioxide; Tissue engineering; Biodegradable polymers; Scaffolds; Growth factor release

1. Introduction Tissue engineering seeks to repair or replace damaged tissues using a combination of cell/molecular biology and materials chemistry/engineering. Over the past decade there has been an intense research effort in this area, which has lead to the generation of a portfolio of tissue engineering strategies. These include, (i) direct implantation of isolated or cultured cells, (ii) implantation of tissues which have been pre-cultured in vitro, and (iii) direct in situ tissue regeneration [1,2]. In many applications a scaffold is required to provide the means for cell/tissue delivery to the repair site and to improve and control the environment for cell growth and tissue maturation. Current challenges to this field include the provision of adequate and appropriate cell sources (e.g. autologous vs. allogeneic vs. xenogeneic cells [3], mature vs. adult stem vs. embryonic stem cells [4,5], etc), vascularisation of engineered tissues [1], improved scaffold materials, appropriate bioreactors and economical scale up and process automation [6].

* Corresponding author. Tel.: +44-0-115-951-5104; fax: +44-0-115951-5110. E-mail address: kevin.shakesheff@nottingham.ac.uk (K.M. Shakesheff).

1359-0286/$ - see front matter Ó 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.cossms.2003.12.004

The scaffold environment should be able to present and deliver combinations of delicate biological materials such as cell adhesion motifs, growth factors, angiogenic factors, differentiation factors and immunosuppressive or anti-inflammatory agents [1,2,7,8]. For all these factors, temporal and spatial control of release are crucial to successful tissue repair and regeneration. The delivery of, for example, growth factors from the scaffold should be controlled such that the factors are delivered in appropriate concentrations at the required time in the regeneration process, as is the case in natural wound healing processes [8,9]. In addition, any scaffold processing should not alter the activity of any active biological agent, e.g. denaturation of proteins through exposure to heat or organic solvents. A variety of scaffolds have been used to date, including natural or synthetic polymers that generally form either hydrogel or monolithic solid polymer scaffolds [1,3,10]. Injectable hydrogel systems are particularly attractive for non-invasive approaches to in situ tissue repair but often lack the structural stability of solid monolithic scaffolds. However, solid monolithic porous scaffolds are frequently difficult to effectively seed with cells due to their often tortuous matrices, and will require surgically-invasive implantation. Natural polymer scaffolds, including hyaluronic acid (HyA) [11,12], collagen [13,14], alginate [15] and fibrin [16], have all been successfully used for various tissue

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Table 1 Summary of advantages and disadvantages of current scaffold fabrication methods Fabrication method

Advantages

Disadvantages

Fibre bonding [27]

Improved stability over non-bonded tassels and felts

Solvent casting/particulate leaching [28]

Control over porosity and pore size ( 6 93%)

Melt moulding/particulate leaching [29]

Control over porosity and pore size ( 6 93%); no organic solvents High volume of interconnected porosity ( 6 97%) High volume of interconnected porosity ( 6 97%) High degree of spatial control of scaffold architecture High degree of spatial control of scaffold architecture

Limited application as the two polymers must be immiscible in solvent and melt state; no real control over porosity and pore size; poor mechanical properties Only produce thin membranes; use organic solvents High temperature required

Emulsion freeze drying [30] Phase separation [31] 3D printing [32] Fused deposition modelling [33]

targets. Natural scaffolds can be processed to retain much of the biological information they contain, or modified by enzymatic means to include a variety of biological agents [17,18]. Degradation of natural scaffolds is controlled by enzyme release from cells within the matrix. These materials can, however, suffer from problems such as poor mechanical strength, interbatch variation, immunogenicity and potential disease transmission [19,20]. Synthetic polymers, including poly(ethylene glycol) (PEG), poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactide-co-glycolide) (PLGA), poly(caprolactone) (PCL) and poly(propylene fumarate), have also been used extensively as tissue engineering scaffolds [3,10,21,22]. Resorption of poly (a-hydroxy acids) is controlled by random hydrolysis and occurs as a bulk degradation [23]. This degradation is exploited to provide a controlled release mechanism for encapsulated therapeutics, and ensures that there is no long-term foreign implant. Unfortunately, acidic by-products from the breakdown of these polymers can lead to inflammatory reactions in some circumstances [24]. However, their FDA approval and long history of use mean they remain an enduring presence in many tissue engineering strategies. Early research efforts in tissue engineering with synthetic scaffolds used degradable scaffolds made from non-woven surgical meshes. Non-woven meshes are suited to the culture of certain tissue types, e.g. cartilage [25,26], but lack the structural stability for many applications. A plethora of new fabrication methods have therefore since been investigated, with the choice and design of the scaffold and processing method depending on a number of factors. Key requirements for scaffold fabrication methods include control of porosity/pore size, maintenance of mechanical properties, and maintenance of material biocompatibility. Other desirable features include avoiding the use of organic solvents, or if they are required, complete removal of solvent residues after processing.

User/equipment sensitive; use of organic solvents User/equipment sensitive; use of organic solvents Use of organic solvent with PLA/PGA Polymer heated to melt state

The ability to incorporate delicate bioactives without the denaturation that may occur upon exposure to solvents, high temperature or shear stresses is also important. Table 1 summarises some of the current fabrication methods. Each method has its own particular advantages and disadvantages, with the choice of method often depending upon the end application. With the drawbacks to these methods in mind, this review will focus upon the emergence of supercritical fluid (SCF) technologies as an alternative or complimentary processing method.

2. Supercritical fluid technologies A supercritical fluid (SCF) is created once a substance is exposed to an environment where its critical temperature and pressure (the intercept of which is referred to as the critical point) are exceeded. Under these conditions, the liquid and gaseous components of the material become identical (Fig. 1), and further compression of this fluid phase will not result in liquefication. SCFs combine the properties of the two phases from which they are formed; they have densities and solvating properties similar to those of liquids, but in combination with a diffusivity and viscosity comparable with that of a gas. The solvent properties of these materials can in fact be controlled very precisely by small manipulations in the pressure (and therefore density) at with the SCF is used. CO2 is the most common candidate for use as a SCF due to its low toxicity, flammability and cost, ready availability, stability, and environmental acceptability. In addition, the critical point conditions of 31 °C and 73.8 bar are readily attainable. As such, supercritical CO2 (scCO2 ) has been employed in a diverse range of applications, including polymer synthesis [34,35], drug delivery [36,37], powder production (e.g. proteins [38,39] and ceramics [40]), powder coating

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Fig. 1. Chronological images of a view cell showing the creation of the single supercritical fluid phase (image c) by taking liquid CO2 above its critical temperature and pressure (images a and b); with subsequent lowering temperature (and hence pressure) the process is reversed (images d and e).

[41], dyeing [42], impregnation [43], and lithography [44]. Exploitation of the unique properties of SCFs has led to the development of a number of polymer processing strategies. There are many excellent reviews outlining these processes in more detail [36,45], so, as this review is concerned primarily with the processing of tissue engineering scaffolds, the brief discussion of these methods will be limited to manipulation of such materials. As such, there are three main types of processing strategy in use: 1. rapid expansion of supercritical solutions (RESS), 2. particles from gas saturated solutions (PGSS), 3. antisolvent techniques. 2.1. RESS technique In polymer processing, this method is often used to generate microparticles for drug delivery applications [46,47]. Here, the polymer and drug must dissolve in the SCF. This mixture is then rapidly expanded into a low temperature and pressure environment. This leads to a rapid decline in the solubility of the polymer in the SCF, and crystallisation of the solute as micro- or nanoparticles of a narrow size distribution. RESS can be used for any polymer with sufficient solubility in the SCF. Unfortunately, most polymers and pharmaceuticals have a very low solubility in scCO2 , except for their low molecular weight fractions. Despite these drawbacks, this approach has already been used to generate PDL LA microparticles containing a cholesterol-lowering statin agent [48]. 2.2. PGSS technique The effect of scCO2 on many polymers is to lower their glass transition temperatures (Tg ) and, if the Tg is reduced below the operating temperature, the material will plasticize, or liquefy. This is because of the very high solubility of the dense gas in the polymer. Some poly-

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mers are in fact able to absorb large concentrations (10–40 wt%) of CO2 [45], which enables the polymer to liquefy at a temperature well below its Tg . The SCF is therefore applied under pressure into the polymer until a gas saturated solution is formed. When depressurized through a nozzle, the gas comes out of liquefied polymer, and the polymer foams, or forms into well defined particulates. This process has the advantages that the starting material does not have to be soluble in the SCF, and that no organic solvents are required during processing. This technique has been previously employed to fabricate microparticles of nifedipine. In this example, the drug molecule itself was shown to plasticize in the presence of the SCF, thus leading to production of fine particles with enhanced release properties [49]. Additionally, a mixture of PEG and nifedipine was found to micronise to particles. A variant of the process developed by Shine and Gelb has also been developed to microencapsulate viral materials within polymers. The technique was termed polymer liquefication using SC solvation (PLUSS) [45]. 2.3. Antisolvent techniques The Gas Antisolvent (GAS) technique [36,50] employs a SCF to act as an antisolvent and precipitate a solute from an organic solvent. As the liquid phase is soluble with the SCF, it expands upon its addition, which results in a decreased capacity of the solvent to support polymer dissolution. This results in the precipitation of microparticles, the size of which can be controlled by adjusting the temperature, pressure or rate of gas addition. This strategy has been used to prepare PL LA [51] and PEG/PLA blend nanoparticles [52] for incorporation and delivery of insulin, with activity of the protein being shown to be maintained during processing. Other related antisolvent techniques include precipitation with compressed antisolvents (PCA) [53,54], supercritical antisolvent (SAS) processing [55], aerosol spray extraction system (ASES) [56], and solution enhanced dispersion by SCFs (SEDS) [57].

3. Uses of SCFs in tissue engineering Supercritical fluids have been applied in recent years to developing tissue engineering strategies. The impetus for embracing these processing methodologies has ranged from their potential ability to create scaffolds with controlled porosities [58], the creation of encapsulated growth factors [59,60], pharmaceuticals [36] and plasmids [61] for controlled delivery to developing tissues, the removal of residual solvents following other fabrication methods [62], and even the treatment of tissue samples to improve their biointegration (e.g. bone delipidation to reduce the immunogenicity of allogeneic

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grafts [63,64]). Due to the ambient processing temperatures, and the potential to process polymers without employing organic solvents, the supercritical approach is also attractive for the potential to preserve the activity of delicate protein molecules that may form an integral component of the tissue regeneration process. Tissue engineering scaffolds are required to possess both a high degree of porosity and an interconnected pore structure. This ensures adequate cell attachment throughout the 3D matrix, promotes both rapid angiogenesis throughout the construct and integration of the implant with host tissue, and facilitates cell migration and nutrient transfer during tissue morphogenesis. Porosity is created during the SCF process by the thermodynamic instability that results following depressurisation around the liquefied polymer [65,66]. This results in the nucleation of the gas molecules in order to minimize their free energy, which creates pores as the polymer resolidifies. The porosity and pore structure of the resultant scaffold can therefore be controlled by varying the amount of gas incorporated and, as shown in Fig. 2, its release rate from the polymer (which is in turn related to the pressure and temperature to which the polymer is exposed and saturated with CO2 ), or the gas diffusion rate (controlled by temperature and depressurisation rate). The nucleation of pores will occur both homo- and heterogeneously [58,67], that is, will form either throughout the continuous polymer phase, or at an interface within the material. Heterogeneous nucleation occurs more readily than the homogeneous route due to the work and free energy required in creating a volume change and forming a stable new surface. Factors such as the inherent porosity in the preprocessed material and residual solvents that may plasticize and alter the surface free energy of the polymer can therefore impact upon the extent of each nucleation type, the differing rates of which inversely affect the average pore size formed [58].

Fig. 2. SEM micrographs of PD; L LA polymer after processing in scCO2 at 240 bar, 35 °C with (a,b) 12 min venting and (c,d) 60 min venting.

In 1990, De Ponti et al. described a Ôgas foaming’ process for fabricating scaffolds from the biodegradable poly(a-hydroxyacid)s (e.g. PL LA, PDL LA, PGA and PLGA) commonly employed in tissue engineering research [68]. Subsequently, Mooney et al. [58] exposed either compressed polymer pellets or solvent cast discs to CO2 at a pressure of 55 bar at 20–23 °C for 72 h (see Table 2). Although not supercritical, over prolonged periods, CO2 under these conditions is still able to saturate the polymer and create a gas saturated polymer, from which depressurisation/gas nucleation can result in a vitrified, porous scaffold being formed. It was confirmed that a high amorphous fraction is a prerequisite for this processing technique as, of the materials tested, crystalline polymers (PGA, PL LA) did not form porous structures due to their relative inability to dissolve gases. The intrinsic viscosity of the polymer must also be considered, as longer polymer chains, being more entangled, display an increased resistance to expansion during nucleation, and thus result in a smaller pore size [72]. Two potential drawbacks that were noted with this strategy were the formation of a non-porous skin on most samples (which results from rapid diffusion of gas away from the surface [58,85]), and a lack of interconnectivity between pores (limited to between 10% and 30% of the total pore number [21]). These problems were successfully addressed by the introduction of NaCl particles as a porogen prior to CO2 exposure, which is leached out afterwards [69]. As this process relied on the close proximity of randomly distributed porogen, further control over the degree of interconnection was later achieved by fusion of adjacent salt particles following exposure to a high humidity environment [71]. When using this strategy, it has been shown that consideration should be given to the ratio of salt to polymer [75] and the gas exposure time used [72], in order to ensure that the scaffolds created in this manner possess suitable mechanical strength. This method of scaffold fabrication has been employed both for the controlled delivery of proteins and as a support for the 3D culture of numerous cell types (Table 2). Of particular importance has been the demonstrated dual delivery of two angiogenic growth factors (vascular endothelial growth factor; VEGF, and platelet-derived growth factor; PDGF), each with their own tailored release profiles. This was achieved by simultaneously processing protein-encapsulated polymer microparticles and a powder/protein mix using this Ôgas foaming’ approach [74]. The microparticle-encased factor is dispersed evenly throughout the resultant scaffold matrix and is released as the polymer degrades, whereas the factor that is mixed freely is associated largely with the surface of the polymer and is therefore released more rapidly. VEGF is an initiator of angiogenesis while PDGF has been shown to promote maturation of blood

Table 2 Summary of some SCF processes reported for the fabrication of tissue engineering scaffolds Scaffold material

Pre-process presentation

CO2 process conditions

Scaffold characterisation

Pressure (bar)

Temp (°C)

Exposure time (h)

Pr. release rate (min)

Porosity (%)

Ave pore size (lm)

Demonstrated application (if any)

Comments

Ref.

55

20–23

72


0.25

PLGA ¼ 97 PL LA ¼ 72 PDL LA ¼ 67

10–100

Skin layer on most sample surfaces

[58]

PLGA (50:50) PGA

Heat compression moulded discs

55

20–23

72


0.25

PLGA ¼ 94 PGA ¼ 0

100–500 (PLGA)

PGA can be mixed with PLGA to regulate porosity Skin layers on sample surfaces

[58]

PLGA (85:15)

Polymer pellets mixed with NaCl

55

Room

48

–

85–96.5

193–439

Smooth muscle cell culture

No skin present Interconnected pores Salt fusion further enhances interconnectivity

[69–71]

PLGA (85:15) + alginate

Compression moulded with NaCl

59

Room

20

2

93

–

Mineralized for bone engineering; VEGF delivery

Scaffolds mineralized by incubation in simulated body fluid VEGF over 70% active for up to 12 days. 44% incorp efficiency––losses during leaching/mineralization

[59]

PDL LA

Compression moulded with and w/o NaCl and alginate

59

Room

Up to 48

1–10

PLGAs ¼ 95 PDL LA ¼ 0


360

VEGF delivery

Very rapid pressure release reduces porosity by 2% Short exposure times (<6 h) creates fragile scaffolds VEGF over 100% active compared with controls, with incorp efficiency 90% and 72% for PLGA 85:15 and 75:25 respectively

[72,73]

PLGA (75:25; 85:15)

Double emulsion microspheres and particles with alginate

55

Room

72


0.25

–

–

Dual delivery of PDGF and VEGF; angiogenesis

Release profiles controlled by the respective growth factor distributions throughout polymer

[74]

PLGA (75:25)

Double emulsion microspheres and NaCl

55

Room

12–16

Rapid

>94

–

DNA delivery

Salt:polymer ratio between 25– 29:1 required to give interconnectivity and mechanical strength

[61,75]

PLGA (80:20; 65:35)

Emulsion with aqueous protein phase

80

35

24

0.17–0.2

–

–

bFGF delivery

Residual solvent above acceptable USP levels––further solvent removal required

[76]

PDL LA PLGA (75:25) PCL

Polymer powders

172

37

<1

2–120

–

100–500 and 0.05–5

Hydroxyapatite composites; protein delivery (inc. BMP-2 and pleiotrophin); bone engineering; liver

One-step process Interconnected pores created without need for porogens Enzyme activity 100% retained

[60,77–82]

PLGA (50:50; 75:25 and 85:15) PGA

PGA ¼ 0

317

Solvent cast discs

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PDL LA (50:50) PDL LA PL LA

Antisolvent system Various pore sizes due to nanofilament structure and CO2 washing steps

[83] Process for improved dispersion of low protein loadings High pressure minimizes skin formation

a

As measured by mercury porosimetry, which only measures the pore volume of interconnected pores [70].

0.5–1000 – 0.66+ Dissolved in DMSO Hyaluronic acidbased polymers

90 (120 for washing steps)

36

–

10–600 60–70a 0.33 Lyophilized powder mixed with protein PDL LA

207

35

8

Ave pore size (lm) Porosity (%) Temp (°C)

Exposure time (h)

Pr. release rate (min) Pressure (bar)

Protein delivery

Comments Demonstrated application (if any) Scaffold characterisation CO2 process conditions

Pre-process presentation Scaffold material

Table 2 (continued)

[84]

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Ref.

318

vessel networks by recruitment of smooth muscle cells. Dual delivery of these two factors each, with distinct release kinetics, promoted the formation of a vascular network in vivo to a greater extent than either factor delivered alone. Such complex release kinetics will prove essential in order to mimic the intricate cascades of growth factors that are stimulated at precise stages during tissue formation and repair [8,86]. It should be pointed out that the drawback of the salt leaching route to controlled porosity is that the leaching stage does remove some of the incorporated bioactive material. Approaches for poly(a-hydroxyacid) scaffold production using scCO2 have been reported by Hile et al. [76]. Here, a water-in-oil emulsion, consisting of an aqueous protein (basic fibroblast growth factor; bFGF) phase and an organic polymer solution phase, is exposed to a SCF in order to (i) extract the organic solvent and precipitate the polymer, (ii) saturate the polymer phase and then, upon pressure release, (iii) create a porous structure. Reported residual solvent levels were, however, above US Pharmacopoeia levels, and further solvent removal is therefore required before these constructs would be suitable for in vivo use. The method of Howdle et al. [61,87] is based upon the PGSS approach and completely avoids the use of solvents. Scaffolds such as those shown in Fig. 3 are instead produced in a one-step process, a protein being first homogeneously distributed throughout the plasticized polymer by use of an impeller device [83,88]. The rate of depressurisation has been shown to affect the pore size distribution of the resultant matrices, with a faster depressurisation (2 min) creating smaller pores. The process also leads to the generation of micropores throughout the scaffold i.e. interconnectivity is created without the use of porogens. This approach has been used to create PDL LA biocomposites containing high loadings of calcium hydroxyapatite [87]. The enzyme ribonuclease A was also incorporated and released over an 80 day period, with activity being completely retained following exposure to the processing conditions [61].

Fig. 3. Fabrication of cylindrical porous PD; L LA scaffolds by SCF processing of the polymer within a shaped mould.

R.A. Quirk et al. / Current Opinion in Solid State and Materials Science 8 (2004) 313–321

Successful bone formation has been demonstrated following the culture of osteoprogenitor cells on these scaffolds in both in vitro and in vivo models using a range of both adhesion peptide [77] and growth factor modifications [78–80]. An alternative to the creation of solid scaffolds is to force the polymer through an orifice during depressurisation and to thus fabricate microparticles [89]. Microparticle-based scaffolds are finding application as injectable systems for cell and growth factor delivery within the body direct to a site of injury or repair [90]. Despite poly(a-hydroxyacid)s remaining the most popular choice of scaffold material (and therefore featuring heavily in SCF fabrication strategies), other polymers are also beginning to be processed into 3D matrices using this route. These include the crystallisation of swollen, crosslinked low density poly(ethylene) using SC propane to create porous structures for the culture of hepatocyte HepG2 cells [91], and the production of solid structures using hyaluronic acid-derivatives [84]. HyA is a naturally-occurring glycosaminoglycan, and is increasing in popularity as a bioresorbable scaffold material. Its use to date has been limited by its high solubility in aqueous solutions at body temperature; an effect that can be remedied by the creation of polymers of HyA esters [92]. These materials have been used in the areas of bone [93] and liver [94] regeneration, and in urethral reconstruction [95]. Employing an antisolvent (SAS) SCF process, these polymers can be fabricated into sponges, threads, and microparticles, with negligible residual solvent being evident after processing.

4. Summary In recent years, SCFs have been employed to address the drawbacks of existing scaffold fabrication methods. Key advantages are to avoid the use of conventional solvents or, if required, to eliminate organic solvent residues. Additionally, they can be used to create biocomposites and/or preserve protein activity, and to exert a high level of control over scaffold porosity and architecture. Numerous processing strategies are emerging which rely on SCFs for either their plasticizing or antisolvent properties. However, despite the obvious advantages, there are some limitations. The control over internal scaffold architecture cannot approach that of the 3D printing technique, and the range of polymer types for which SCFs are applicable might be limiting in some applications, particularly where high mechanical strength is required. Further process optimisation is therefore still required, but the future of SCFs looks bright, and the versatility of SCF processing will ultimately result in innovative combinations with other scaffold fabrication methodologies.

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Acknowledgements We thank JJA Barry and MMCG Silva for their help and advice. SMH is a Royal Society-Wolfson Research Merit Award Holder.

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